Wired implantable monolithic integrated sensor circuit

ABSTRACT

There is provided a glucose sensor system comprising: a transmitter (2) for containing a battery (212), the transmitter being for placement on top of patient skin; a transcutaneous connector (3) comprising at least one conductive path; and an implantable monolithic integrated circuit (I) for placement beneath the patient skin, wherein the implantable monolithic integrated circuit comprises a potentiostat and an electrochemical sensing element; wherein the potentiostat is electrically coupled to the transmitter (2) via the transcutaneous connector (3), and the electrochemical sensing element is configured to sense glucose concentration and generate an electrical signal representative of the glucose concentration, and wherein the potentiostat is electrically connected to the electrochemical sensing element.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims any and all benefits as provided by lawincluding benefits under 35 U.S.C. 119(e) of U.S. ProvisionalApplication No. 63/239,484, filed Sep. 1, 2021, U.S. ProvisionalApplication No. 63/298,632, filed on Jan. 12, 2022, U.S. ProvisionalApplication No. 63/336,299, filed on Apr. 28, 2022, U.S. ProvisionalApplication No. 63/318,790, filed on Mar. 11, 2022, and U.S. ProvisionalApplication No. 63/333,443, filed on Apr. 21, 2022, the contents of eachof the above are incorporated herein by reference in their entirety.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

This invention was made with government support under contract no.R44DK111001, and no. R43DK121621 awarded by the National Institutes ofHealth. The government has certain rights in the invention.

BACKGROUND Field of the Invention

The present disclosure provides a scalable, flexible, electrochemicalsensing platform including a transmitter, a wire, and a monolithicintegrated sensor circuit containing one or more integrated sensors.

Description of the Related Art

Continuous monitoring of personal health has the potential torevolutionize healthcare by enabling preventative health managementcompared to traditional treatment strategies which rely upon a fewmeasurements at discrete points. Continuous monitoring should enable thedetection of diseases and conditions before users would be drawn to adoctor's office for a discrete blood test.

Realization of the potential of continuous monitoring of health requiresnew tools to measure analytes in the body effectively and continuously.There are currently some health monitoring devices being used tocontinuously measure in-vivo analyte levels. However, the technologieswhich have been thus far deployed have suffered from shortcomingspreventing widespread adoption.

An excellent example of a continuous monitoring system that has come tomarket can be seen with Continuous Glucose Monitoring (CGM) systems. CGMsystems from MEDTRONIC® and DEXCOM® are available for continuous glucosemonitoring for diabetics using transdermal systems. Less than 10% ofdiabetics currently use Continuous Glucose Monitoring (CGM) systems ofany kind, although it has been shown to be the best method for diabetesmanagement.

The disadvantages of transdermal CGM technologies from, for example,MEDTRONIC® and DEXCOM®, include their high complexity, large size andhigh cost. These devices are prone to rejection by the body due to theimmune system response, use bulky external transmitters for sensorcontrol and data processing, and have high manufacturing costsassociated with utilizing and integrating discrete components. The bulkyexternal transmitters require strong skin adhesives which causeirritation and skin allergies. These devices are also short-lived andrequire frequent removal and insertions. Their response speed andaccuracy can become limiting factors in some situations, e.g., duringsleep, at mealtimes.

In view of these shortcomings, efforts have been made to make better,more affordable, and user-friendly CGM devices.

Implantable CGM systems avoid the transdermal wire set-up of classicalCGM systems and thus reduce infection and skin irritation issues.Implantable glucose monitoring systems commercially available includethose by, for example, SENSEONICS®, including the EVERSENSE® system.However, these implantable systems have suffered from miniscule uptakeby consumers because they are beset by their own limitations.Implantation of the device requires the scheduling of a medicalprocedure to ensure that the device is properly positioned within theuser at great cost. Additionally, removal of the device can bechallenging. First the device must be exactly located. Then a skilledmedical professional must remove the device. Additionally, the risksposed by device movement within the human body have forced implantabledevices to have a large surface area. The exposed area of theimplantable devices has increased the rate of foreign body response andrejection by users.

SUMMARY

Accordingly, a clear need exists for an improved platform in thecontinuous health monitoring industry. The present disclosure addressesseveral of these shortcomings of the prior art to finally makecontinuous health monitoring more mainstream.

In various instances, the present disclosure provides for a monolithicintegrated sensor circuit for placement inside the body which senses theconcentration of an analyte by an electrode and transmits digital datato a transmitter outside the body. The monolithic integrated sensorcircuit and the transmitter are connected by a transdermal wire. Bycombining in a single monolithic integrated circuit, the sensorchemistry driving analyte detection as well as the circuitry needed totranslate the concentration of the analyte into digital information,embodiments of the present disclosure reduce complexity, size, and costof traditional transdermal CGM systems. Additionally, by placing theelectronics adjacent to the electrodes which measure analyte, thepresent disclosure permits the determination of a superior signal for alonger period as compared to traditional systems because of thepermissive use of multiple discrete electrodes with minimal electronicsignal travel. Moreover, this also reduces the complexity of thetransmitter and hence enables a smaller, and lighter transmitter. Thisincreases sensor wear time and decreases the need of strong adhesives,hence minimizing skin irritation and allergies.

Unlike implantable wireless CGM systems, such as the EVERSENSE® system,the embodiments of the present disclosure present a wired connection forthe transdermal delivery of power and, optionally, the exchange ofdigital information. The wired connection permits easier insertion ofthe device as well as a more stable supply of power. The transdermalwired connection also fixes the device allowing a smaller size for thedevice. Lastly, the wire connection allows the device to be much moreeasily removed by the user.

According to a first aspect of the present disclosure, there is provideda transdermal analyte concentration measurement system comprising: 1) animplantable monolithic integrated circuit containing: an integratedsensing element which senses one or more analytes and generates a signalrepresentative of an analyte concentration; an integrated sensor signalacquisition unit which receives and processes the signal from theintegrated sensing element; an integrated communication unit connectedto the integrated sensor signal acquisition unit which transmits datarepresentative of said analyte concentration; and an integrated powermanagement unit connected to and providing power to the sensing element,the sensor signal acquisition unit, and the communications unit; and 2)a transmitter, configured for operation outside of the body of a user,connected by a wire to the integrated communication unit and integratedpower management unit of the implantable monolithic integrated circuit,said transmitter providing power to the integrated power management unitof the implantable monolithic integrated circuit via the wire andreceiving data from the integrated communication unit of the implantablemonolithic integrated circuit via the wire, and said wire configured tocross from the skin surface into the user.

The integrated sensing element of the implantable monolithic integratedcircuit may include one or more integrated electrodes.

At least one of the one or more integrated electrode(s) may comprise aconductive surface of one or more conductive materials.

The integrated electrode(s) may be made using lithography.

The integrated electrode(s) may be coated with a hydrogel formed from across linking agent, an enzyme, and a proteinaceous material. Thehydrogel may further comprise one or more co-protein. The co-protein(s)may be configured to improve longevity, decrease foreign body response,and/or increase signal of the integrated sensing element.

The cross linking agent may comprise glutaraldehyde. Further oralternatively, the enzyme may comprise glucose oxidase. Further oralternatively, the proteinaceous material may comprise human serumalbumin.

The hydrogel may further comprise a co-protein of catalase orhorseradish peroxidase.

The wire may be comprised within a flexible printed circuit board.

The transdermal analyte concentration measurement system may comprisethe flexible printed circuit board. The wire may comprise more than oneconductive path. In particular, the wire may comprise at least twoconductive paths.

According to a second aspect of the present disclosure, there isprovided a glucose sensor system comprising: a transmitter forcontaining a battery, the transmitter being for placement on top ofpatient skin; a transcutaneous connector comprising at least oneconductive path; and an implantable monolithic integrated circuit forplacement beneath the patient skin, wherein the implantable monolithicintegrated circuit comprises a potentiostat and an electrochemicalsensing element; wherein the potentiostat is electrically coupled to thetransmitter via the transcutaneous connector, and the electrochemicalsensing element is configured to sense glucose concentration andgenerate an electrical signal representative of the glucoseconcentration, and wherein the potentiostat is electrically connected tothe electrochemical sensing element.

It would be understood that a monolithic integrated circuit (alsoreferred to as an IC, a chip, or a microchip) is a set of electroniccircuits integrated on one small piece (or chip) of semiconductormaterial (for example but not limited to silicon).

By combining the electrochemical sensing element and the potentiostat ina single monolithic integrated circuit, the glucose sensor system has animproved performance and reduced complexity, size and costs. Inparticular, the short distance (e.g., in microns) between theelectrochemical sensing element and the potentiostat allows for minimalelectronic signal travel. Thus, the glucose sensor system can determinea more superior signal for a longer period, as compared to traditionalsystems where the potentiostat is placed within the transmitter on topof patient skin. This arrangement also enables a smaller and lightertransmitter, and accordingly provides a more pleasant user wearexperience.

It would further be understood that the potentiostat is an electronichardware used for controlling and measuring the electrochemical sensingelement. The potentiostat may comprise an electric circuit whichcontrols (or maintains) the voltage potential of an electrode (e.g., aworking electrode) of the electrochemical sensing element and senseschanges in the resistance of the electrode, by outputting a currentwhich is inversely proportional to the varying resistance of theelectrode. Therefore, the electrochemical sensing element and thepotentiostat collectively generate an electrical signal (e.g., anelectric current) representative of the glucose concentration.

The transcutaneous connector wire may be configured to cross from theskin surface into the user. The glucose sensor system may furthercomprise a reader wirelessly linked to the transmitter. The wirelesstechnology used is one that is suitable for such link. For example,Bluetooth technology is suitable for short distance (e.g., 20 meters)communication. Zigbee is a different technology that can be used for aslightly longer-range (e.g., 100 meters) communication.

The glucose sensor system may further comprise a smart insulin pen orpump wirelessly linked to the transmitter. Such a system can be used foran open-loop or closed-loop glucose control as an artificial pancreas ora part of an artificial pancreas system.

The glucose sensor system may further comprise a secure databasewirelessly linked to the transmitter. The database can be used tocommunicate the long-term data trends and important events (e.g.,hypoglycemia events) to the care team.

The potentiostat may be continuously powered by the battery.

The transmitter and the sensor many have other features (e.g., atemperature sensor, photodetectors) to detect any changes in theenvironment.

By “continuously powered”, it is meant that when the battery containssufficient electric power (e.g., before it is drained up), the batterycontinuously supplies power to the potentiostat via the transcutaneousconnector. By continuously powering the potentiostat, the potentiostatis able to maintain the potential of the working electrode(s) of theelectrochemical sensing element continuously without any break. In thisway, there is no need to frequently calibrate the glucose sensor systemduring the lifetime of the battery.

The glucose sensor system maybe a continuous glucose sensor system.

The continuous glucose sensor system may also be referred to as aContinuous Glucose Monitoring (CGM) system. A CGM system essentiallymeasures glucose levels 24/7, from once every tens of seconds to onceevery few minutes. The measurement frequency may be adjusted at thetransmitter side. Generally speaking, the CGM system requires thepotential of the working electrode(s) of the electrochemical sensingelement to be continuously maintained.

The transcutaneous connector may be a flexible connector.

The transcutaneous connector may comprise a printed circuit board. Theprinted circuit board may be a flexible printed circuit board (e.g., theflexible printed circuit board according to the ninth aspect describedbelow).

The potentiostat may be placed at a depth of 1 to 10 mm beneath thepatient skin.

The potentiostat may be at a depth of 1 to 5 mm beneath the patientskin. More specifically, the potentiostat may be at a depth of 2 to 3 mmbeneath the patient skin. The 2-3 mm depth allows the potentiostat togenerate a superior signal indicative of glucose concentration, and alsosignificantly shortens the communication distance between thepotentiostat and the transmitter.

The electrochemical sensing element may comprise at least one workingelectrode coated with a chemistry which converts glucose concentrationinto current.

The electrochemical sensing element may be configured to transducechanges in glucose concentration into changes of an electrical current.

The electrical signal representative of the glucose concentration maycomprise the current.

The potentiostat may be configured to control a potential of the atleast one working electrode based upon a reference voltage (e.g.,V_(WE)). The potential of the at least one working electrode iscontrolled to be substantially the same as the reference voltage.

The reference voltage may be programmable by the transmitter.

The implantable monolithic integrated circuit may further comprise adigital to analog converter configured to generate the reference voltagebased upon a user input transmitted from the transmitter via thetranscutaneous connector.

The potentiostat may be further configured to buffer the current to ananalog to digital converter connected to the potentiostat.

The analog to digital converter may be configured to convert the currentinto a digital signal.

The electrochemical sensing element may further comprise a referenceelectrode and a counter electrode. The counter electrode may be used tobalance the current generated by the at least one working electrode. Thereference electrode may be used to provide a stable voltage referencesignal beneath the patient skin.

The at least one working electrode may share a common referenceelectrode.

The at least one working electrode may share a common counter electrode.Alternatively, the at least one working electrode may each be associatedwith a respective counter electrode.

The potentiostat may be configured to maintain a predetermined voltagebetween the at least one working electrode and the reference electrodewhile providing current through the counter electrode.

The potentiostat may comprise a first operational amplifier controllinga voltage of the reference electrode (e.g., through negative feedback)while providing current through the counter electrode and a secondoperational amplifier controlling a voltage of the at least one workingelectrode (e.g., through negative feedback) and buffering the current ofthe at least one working electrode to an analog to digital converterconnected to the potentiostat. Advantageously, this arrangement of thepotentiostat allows for independently controlling the potentialdifference between working and reference electrodes in a multi-analytesensor where there are multiple working electrodes for sensing differentanalyte.

The at least one working electrode may comprise a conductive surface ofone or more conductive materials. The conductive material can be a noblemetal. The conductive metal can be platinum. The at least one workingelectrode may be made using lithography.

The chemistry may comprise a hydrogel including a cross linking agent,an enzyme, and a proteinaceous material. The hydrogel may furthercomprise a co-protein. The co-protein may be configured to improvelongevity, decrease foreign body response, and/or increase signal of theintegrated sensing element.

The cross-linking agent may comprise glutaraldehyde. The enzyme maycomprise glucose oxidase. The proteinaceous material may comprise humanserum albumin.

The hydrogel may further comprise a co-protein of catalase orhorseradish peroxidase.

The potentiostat may be connected within 2 millimeters to the at leastone working electrode. More preferably, the potentiostat may beconnected within half a millimeter to the entirety of the at least oneworking electrode. By the expression “the entirety of the at least oneworking electrode”, it is meant that the potentiostat may have differentconnection distances to different ones of the working electrodes, butthe distance between the potentiostat and the furthest working electrodeis within half a millimeter.

The at least one working electrode may be patterned to increase surfacearea. Utilizing patterned or non-planar working electrodes (instead ofconventional planar electrodes) increases the effective signal strength,because the strength of the sensing element signal can be proportionalto surface area of the electrode.

The at least one working electrode may be patterned by forming pillars.The pillars may have a spacing of 0.25 μm-25 μm and a height of 0.1μm-10 μm.

The at least one working electrode may be integrated with thepotentiostat into the same CMOS die.

The implantable monolithic integrated circuit may be CMOS based. Theimplantable monolithic integrated circuit and the CMOS die may be usedinterchangeably.

The implantable monolithic integrated circuit may be from 30 microns to600 microns in thickness. In particular, the implantable monolithicintegrated circuit may be from 50 microns to 150 microns in thickness.

The implantable monolithic integrated circuit may be from 30 microns to600 microns in thickness, 500 microns to 10,000 microns in length and ina range from 100 microns to 4,000 microns in width. More preferably, theimplantable monolithic integrated circuit may be from 50 microns to 150microns in thickness, 1,500 microns to 3,000 microns in length and in arange from 100 microns to 4,000 microns in width.

The implantable monolithic integrated circuit may comprise holesextending therethrough.

The implantable monolithic integrated circuit may be coated with abiocompatible soft material.

The implantable monolithic integrated circuit may comprise roundededges.

The implantable monolithic integrated circuit may comprise a sensorsignal acquisition unit which receives and processes the electricalsignal from the electrochemical sensing element. The sensor signalacquisition unit may comprise the potentiostat.

The implantable monolithic integrated circuit may further comprise acommunication unit electrically connected to the sensor signalacquisition unit and transmitting data representative of the glucoseconcentration.

The data representative of the glucose concentration may be based uponthe processed electrical signal by the sensor signal acquisition unit.

The implantable monolithic integrated circuit may further comprise apower management unit electrically connected to and providing power tothe electrochemical sensing element, the sensor signal acquisition unit,and the communication unit.

The transmitter may be electrically connected by the transcutaneousconnector to the communication unit. The transmitter may be configuredto receive data from the communication unit via the transcutaneousconnector.

The data received by the transmitter may comprise the datarepresentative of the glucose concentration, and optionally temperaturedata. The temperature data may be generated by a temperature sensor ofthe implantable monolithic integrated circuit.

The transmitter may be configured to transmit data to the communicationunit via the transcutaneous connector.

The transmitter may be configured to communicate with the communicationunit in a bidirectional and time multiplexed fashion.

The communication unit may comprise a receiver subsystem and atransmission subsystem.

The transmitter may be electrically connected by the transcutaneousconnector to the power management unit, and the transmitter may beconfigured to provide power to the power management unit via thetranscutaneous connector.

The power management unit may comprise a voltage regulator which isconfigured to regulate the power from the transmitter into a DC voltage.

The power management unit may further comprise a reference generatorwhich is configured to generate reference voltage and/or currents usedby the sensor signal acquisition unit (e.g., the potentiostat, theanalog to digital converter).

The communication unit may further comprise a MUX/DEMUX (i.e.,couple-decouple) network which is configured to decouple the data fromthe power.

The electrochemical sensing element may comprise at least three workingelectrodes with each connected to a respective potentiostat. The use ofat least three working electrodes is beneficial for reducing noise inthe electrical signal representative of the glucose concentration andfor improving the accuracy of the glucose sensor system. In particular,the output of the at least three working electrodes may be compared withone another and further processed (e.g., by averaging and/or removingone output which significantly deviates from the other output) to reducethe impact of noise or fault occurring in some of the electrodes. Theglucose sensor system may further comprise an analog to digitalconverter for placement beneath the patient skin, wherein the analog todigital converter is electrically connected to the potentiostat.

The implantable monolithic integrated circuit (in particular, the sensorsignal acquisition unit thereof) may comprise the analog to digitalconverter.

The glucose sensor system may further comprise a multiplexerelectrically connected between the potentiostats associated with the atleast three working electrodes and the analog to digital converter. Themultiplexer allows the analog to digital converter to be shared by theat least three working electrodes through time division multiplexing.

The implantable monolithic integrated circuit (in particular, the sensorsignal acquisition unit thereof) may comprise the multiplexer.

The at least three working electrodes may be functionalized towardsglucose.

The glucose sensor system may further comprise a control logicconfigured to denoise the information from one or more of the at leastthree working electrodes.

The control logic may compare and/or further process the informationobtained from the at least three working electrodes, by for exampleaveraging the information and/or discarding information obtained fromone electrode which significantly deviates from information obtainedfrom other electrodes. In this way, the noise level contained in theinformation can be reduced.

The glucose sensor system may further comprise a temperature sensor forplacement beneath the patient skin.

The implantable monolithic integrated circuit may comprise thetemperature sensor.

The temperature sensor may comprise a bandgap circuitry to generatecurrent proportional to absolute temperature (PTAT) and currentcomplementary to absolute temperature (CTAT). The glucose sensor systemmay further comprise a control logic programmed to process informationfrom the at least one working electrode by taking into accountinformation from the temperature sensor.

Advantageously, taking into account information from the temperaturesensor is useful for improving reliability of glucose leveldetermination. The control logic may calibrate the information from theat least one working electrode using the information from thetemperature sensor.

The control logic described above may also be referred to as aprocessor, a processing unit, a controller, or a microcontroller. Theimplantable monolithic integrated circuit may comprise the controllogic. Alternatively, the transmitter may comprise the control logic.Further alternatively, the reader wirelessly linked to the transmittermay comprise the control logic.

The transcutaneous connector may comprise at least two conductive paths,and one of the conductive paths may be a ground connected to thepotentiostat and another one of the conductive paths may function asdata over power. In other words, the another one of the conductive pathsis used to carry both a power signal and a data signal, with the datasignal superimposed over the power signal.

Alternatively, the transcutaneous connector may comprise threeconductive paths, with two paths used to carry power and ground signalsand a third path used for data communication to and from theelectrochemical sensing element.

The implantable monolithic integrated circuit (or the CMOS die) may bebonded to the flexible connector and the at least one conductive path bywire bonding. Alternatively, the implantable monolithic integratedcircuit (or the CMOS die) may be bonded to the flexible connector andthe at least one conductive path by flip chip packaging.

The at least one working electrode may be coated with a glucose oxidasehydrogel.

According to a third aspect of the present disclosure, there is provideda transdermal analyte concentration measurement system comprising: asensing element for placement beneath patient skin and for sensing oneor more analytes, and wherein the sensing element comprises at leastthree working electrodes each of which is configured to generate asignal indicative of a concentration of an analyte; and a control logicconfigured to generate a fused signal representative of the analyteconcentration within the patient based upon at least some of the signalsgenerated by the at least three working electrodes.

The control logic and the sensing element may be an integrated device.The control logic and the sensing element may be integrated within animplantable monolithic integrated circuit.

The transdermal analyte concentration measurement system may furthercomprise a transmitter configured for operation outside of a body of thepatient. The transmitter may be coupled to the sensing element via atranscutaneous connector. The transmitter may comprise the controllogic.

The transdermal analyte concentration measurement system may furthercomprise a device wirelessly coupled to the transmitter. The device maycomprise the control logic.

The control logic may be configured to generate a weighted average ofthe signals generated by the at least three working electrodes. Thefused signal may be the weighted average.

The control logic may be configured to discard one or more of thesignals generated by the at least three working electrodes, and togenerate the fused signal based upon the remaining signals. This meansthat the weight(s) of the discarded signal(s) is zero.

The discarded signal(s) may significantly deviate from the remainingsignals. The fused signal may be generated as a weighted average of theremaining signals.

The discarded signal(s) may significantly deviate from a median value oran average value of the signals generated by the at least three workingelectrodes.

The discarded signal(s), within the signals generated by the at leastthree working electrodes, may have the greatest distance to the medianvalue or the average value. Alternatively, the discarded signal(s) mayhave a distance that is greater than a predetermined distance to themedian value or the averaged value.

The control logic may be configured to filter the signal(s) generated byone or more of the at least three working electrodes so as to removenoise from the signal(s).

The control logic may be configured to generate a series of the fusedsignals at different time points, and to determine a rate of change ofthe analyte concentration within the patient.

The control logic may be configured to alert the patient based upon therate of change.

The control logic may be configured to estimate the future analyteconcentration within the patient based upon the series of the fusedsignals and the rate of change.

According to a fourth aspect of the present disclosure, a continuousglucose monitoring system for a subject is provided comprising: atransmitter for containing a battery, said transmitter for placement ontop of subject skin; a transcutaneous connector comprising at least oneconductive path; an electrochemical sensing element comprising aconductor configured to sense glucose concentration and generate anelectrical signal representative of the glucose concentration; and atemperature sensor located within 10 microns of the electrochemicalsensing element.

Each of the electrochemical sensing element and the temperature sensormay be for placement beneath the subject skin.

By locating the temperature sensor within 10 microns of theelectrochemical sensing element, the temperature sensor is able tomeasure the local temperature at a same location where the glucoseconcentration is sensed. In this way, the temperature sensor can be usedto efficiently calibrate the sensor reading of the electrochemicalsensing element, thereby improving the accuracy of the continuousglucose monitoring system. Further, the temperature reading from thetemperature sensor can be used to accurately indicate the position(e.g., depth) of the electrochemical sensing element (e.g., beneath thesubject skin)

The temperature sensor may not share a conductor with theelectrochemical sensing element.

The temperature sensor may be formed on a monolithic semiconductorsubstrate shared with the electrochemical sensing element. Thesemiconductor substrate may comprise silicon.

The temperature sensor may be a silicon-based bandgap temperaturesensor.

The temperature sensor may contain a pair of semiconductor junctions,and a difference between the voltage across the pair of semiconductorjunctions may indicate a temperature beneath the skin. The semiconductorjunctions may be PN junctions.

The temperature sensor may comprise a pair of bipolar junctiontransistors.

The temperature sensor may be implemented by a circuit that generates acurrent proportional to absolute temperature (PTAT) and a currentcomplementary to absolute temperature (CTAT).

The continuous glucose monitoring system may further comprise a controllogic in communication with the temperature sensor and/or theelectrochemical sensing element. The control logic described above mayalso be referred to as a processor, a processing unit, a controller, ora microcontroller. The control logic may also be formed on themonolithic semiconductor substrate. Alternatively, the transmitter maycomprise the control logic. Further alternatively, a reader wirelesslylinked to the transmitter may comprise the control logic.

The control logic may be configured to process (e.g., calibrate) theelectrical signal representative of the glucose concentration based upona temperature measurement from the temperature sensor, so as to generateprocessed data representative of the glucose concentration.

The control logic may be configured to determine whether theelectrochemical sensing element is inserted beneath the subject skin,based upon a temperature reading from the temperature sensor.

The continuous glucose monitoring system may comprise a second and/orthird temperature sensor. The second and/or third temperature sensor mayshare the same semiconductor substrate as the first temperature sensorand/or the electrochemical sensing element. The control logic may beconfigured to generated a fused temperature data based upon temperaturereadings from two or more of the temperature sensors. The fusedtemperature data may be generated in a way similar to the generation ofthe fused signal according to the third aspect.

The transmitter may comprise a further temperature sensor. The controllogic may be in communication with the further temperature sensor.

The control logic may be configured to determine whether theelectrochemical sensing element is inserted beneath the subject skin,based upon a difference between the temperature readings of thetemperature sensor and the further temperature sensor.

The control logic may be configured to determine a depth of theelectrochemical sensing element beneath the subject skin, based upon adifference between the temperature readings of the temperature sensorand the further temperature sensor.

The control logic may be configured to detect hypoglycemia within thesubject based upon a change of temperature readings from the temperaturesensor and a change of temperature readings from the further temperaturesensor during a same time period. More specifically, if the readingsfrom the temperature sensor indicates a drop of temperature but thereadings from the further temperature sensor during the same time perioddoes not indicate a drop of the same level, it is considered that thesubject is suffering from hypoglycemia. According to a fifth aspect ofthe present disclosure a method for determining the temperature of theepidermis, dermis, and/or subcutaneous tissue of a subject is provided,said method comprising changing the environment of a temperature sensorto a predefined fixed temperature sweep, wherein the temperature sensoris formed on a semiconductor substrate comprising silicon; devising acalibration scheme using raw readings of the temperature sensorgenerated during the predefined fixed temperature sweep, so as to enablethe temperature sensor to accurately determine temperature; attachingthe temperature sensor to a flexible transcutaneous connector comprisingat least one conductive path and a transmitter containing a battery, thetransmitter being for placement on top of the subjects skin; insertingthe temperature sensor into the dermis of the subject while attached tothe transcutaneous connector and transmitter on top of the subjectsskin; and measuring the temperature of the dermis, epidermis and/orsubcutaneous tissue using the devised calibration scheme.

The temperature sweep may comprise varying the temperature between 35 to41 degrees Celsius.

The calibration scheme may be stored in the transmitter.

Devising a calibration scheme may comprise determining calibrationcoefficients, wherein the calibration coefficients are configured to mapthe raw readings of the temperature sensor generated during thepredefined fixed temperature sweep to actual temperatures used bypredefined fixed temperature sweep.

Changing the environment of the temperature sensor comprises placing thetemperature sensor in a saline solution and changing the temperature ofthe salient solution according to the predefined fixed temperaturesweep. According to a sixth aspect of the present disclosure, there isprovided a method for changing the redox potential of an implantedpotentiostat and paired working, control, and reference electrodescomprising: receiving a command for setting the redox potential to apredetermined value, programming (or reconfiguring) a programmablevoltage generator based upon the command, so as to generate a voltagecorresponding to the predetermined value; and feeding the generatedvoltage from the programmable voltage generator to the potentiostat.

It would be understood that the predetermined value is a voltage.

The programmable voltage generator may comprise a digital to analogconverter.

The command may be sent from a transmitter configured for placement onthe skin of a subject and comprising a battery across a flexibletranscutaneous connector including at least one conductive path to theimplanted potentiostat and its paired working, control, and referenceelectrodes.

The digital to analog converter may comprise a resistor ladder connectedbetween a voltage regulator output and a ground to form an m-bit digitalto analog converter capable of generating m different voltages.

The potentiostat and paired working, control, and reference electrodesmay be monolithically integrated on a single semiconductor substratecomprising silicon.

The potentiostat may be configured to control a potential of the workingelectrode based upon the redox potential. The potential of the workingelectrode is controlled to be substantially the same as the redoxpotential.

According to a seventh aspect of the present disclosure, there isprovided a transdermal analyte concentration measurement systemcomprising: a transmitter for placement on top of subject skin; animplantable monolithic integrated circuit for placement beneath thesubject skin, wherein the implantable monolithic integrated circuitcomprises a potentiostat and an electrochemical sensing element forsensing an analyte within the subject, wherein the electrochemicalsensing element comprises at least one working electrode configured togenerate a signal indicative of a concentration of the analyte, and thepotentiostat is electrically connected to the electrochemical sensingelement and configured to control a potential of the at least oneworking electrode based upon a reference voltage; wherein the referencevoltage is programmable by the transmitter.

The transmitter may be configured to send a command to the implantablemonolithic integrated circuit, with the command indicating a desiredvalue of the reference voltage (also referred to as redox potential).

The implantable monolithic integrated circuit may comprises a digital toanalog converter configured to generate the reference voltage based uponthe command.

The electrochemical sensing element may comprise more than one workingelectrode, and the reference voltage of each working electrode may beindependently programmable by the transmitter. Advantageously, thisarrangement allows different working electrodes to have differentreference voltages, thus enabling a wide variety of analytes to bedetected by the transdermal analyte concentration measurement system.According to an eighth aspect of the present disclosure, there isprovided an implantable monolithic CMOS-based integrated circuit forplacement beneath subject skin, comprising a potentiostat and anelectrochemical sensing element for sensing the analyte within thesubject, wherein the electrochemical sensing element comprises at leastone working electrode configured to generate a signal indicative of aconcentration of the analyte, and the potentiostat is electricallyconnected to the electrochemical sensing element, and wherein the atleast one working electrode comprises a surface with a plurality ofholes formed therein.

A bottom surface of the holes may have a combined area which is lessthan an area of the surface of the at least one working electrode. Inother words, the at least one working electrode has a patterned surfacecomprising a hole structure (or a reverse-pillar structure). Utilizingthe patterned or non-planar working electrode (instead of conventionalplanar working electrode) increases the surface area of the workingelectrode, thereby improving the sensitivity of the electrochemicalsensing element. This is because the strength of the sensing elementsignal is generally proportional to a surface area of the electrode. Theholes may be formed (e.g., etched) during a CMOS fabrication process.Further, the reverse-pillar structure is more robust than acomplementary pillar structure.

The surface may be a surface of a metal layer (e.g., a top metal layerof the CMOS fabrication process).

The holes may be formed by etching top metal pillars embedded within topinsulator such that each pillar is microscale in size and has topinsulator removed during CMOS fabrication process such that the hole isexposed for post processing.

The surface may be coated with another metal layer comprising platinum.The at least one working electrode may include interlayer conductivevias. The conductive vias may be formed at a bottom surface of the holesto enable a connection with bottom metals and/or the potentiostat.

The non-planar structure may be coated with a subsequent metal layercomprising platinum.

The holes may be formed by etching top metal pillars embedded within topinsulator such that each pillar is microscale in size and has topinsulator removed during CMOS fabrication process such that the hole isexposed for post processing. The at least one working electrode may becoated with a chemistry which converts the analyte concentration intocurrent. At least a part of the chemistry may fill the holes. Therefore,the patterned surface of the working electrode is also useful forimproving the adhesion between the surface and the chemistry.

The electrochemical sensing element may further comprise a referenceelectrode and a counter electrode. The counter electrode may be used tobalance the current generated by the at least one working electrode. Thereference electrode may be used to provide a stable voltage referencesignal beneath the patient skin. One or more of the counter electrodeand the reference electrode may also comprise a surface with a pluralityof holes etched therein.

According to a ninth aspect of the present disclosure, a flexibleprinted circuit board is provided. The flexible printed circuit boardcomprises a first metal layer configured to route electrical signals, asecond metal layer configured to provide mechanical strength for theflexible printed circuit board without routing any electrical signals,and a dielectric substrate arranged between the first metal layer andthe second metal layer and made of a flexible material.

It would be understood that the first and second metal layers and thedielectric substrate are stacked along a thickness direction of theflexible printed circuit board.

The second metal layer therefore controls the stiffness of the flexibleprinted circuit board. The second metal layer advantageously allows theflexible printed circuit board to have mechanical strength to a certainextent while the entire printed circuit board remains flexible. In thisway, the flexible printed circuit board is suitable for use as atranscutaneous connector in a continuous glucose monitoring system. Thisis because the mechanical strength provided by the second metal layerallows the flexible printed circuit board to be easily inserted under asubject's skin (e.g., by using a needle) and the overall flexibility ofthe printed circuit board is useful for reducing tissue damage withinthe subject in use.

The second metal layer may be made of copper. In same thickness, copperprovides more stiffness than the stiffeners (e.g., FR4, Polyimide,Glass) commonly used in PCBs. Therefore, using copper to control thestiffness of the flexible printed circuit board enables the flexibleprinted circuit board to have a thinner profile as compared to the useof common stiffeners, and hence enable the use of smaller needle sizeswhich reduces insertion pain and foreign body response.

The first metal layer may be a patterned metal layer. The first metallayer may have at least two and at most four conductive traces. Thetraces may range in width and spacing from 1 mil to 5 mils. It would beunderstood that the second metal layer is not one of the conductivetraces.

The flexible printed circuit board may comprise a first insulating layercovering the first metal layer. The first insulating layer and thedielectric substrate may be placed at opposite sides of the first metallayer. The first insulating layer may cover the entirety of the firstmetal layer except at interfacial areas of the first metal layer whichare used to make electrical connections with external components (e.g.,of the continuous glucose monitoring system).

The flexible printed circuit board may comprise a second insulatinglayer covering the second metal layer. The second insulating layer andthe dielectric substrate may be placed at opposite sides of the secondmetal layer. The second metal layer may be completely covered by thesecond insulating layer.

The first and/or second insulating layer may be made from abiocompatible material such as parylene-C or polyimide. The first and/orsecond insulating layer are useful for preventing interaction betweenthe first and/or second metal layers with the body fluids of thesubject.

The first and/or second insulating layer may be made of the samematerial(s) as the dielectric substrate.

The flexible printed circuit board may have a total thickness between 2mils to 20 mils.

The flexible printed circuit board may have a length from 2 mm to 20 mm.

The substrate may be made from polyimide, liquid crystal polymer,polyethylene ethylene, or polyether ether ketone.

The first metal layer may be made of copper. A thickness of the firstmetal layer may be between 0.25 and 0.5 Oz.

A thickness of the second metal layer may be between 0.25 and 2 Oz.

The flexible printed circuit board may be of a dimension which can fitin a 16 gauge to 32 gauge needle.

According to a tenth aspect of the present disclosure, anelectrochemical sensor for measuring interstitial glucose is provided.The electrochemical sensor is configured to detect a rate of changeof >10 mg/dl/minute. The features of “(electrochemical) sensing element”and “integrated sensing element” as set out above also apply to the“electrochemical sensor” of the tenth aspect.

The electrochemical sensor may be configured to detect a rate of changeof >20 mg/dl/minute. The electrochemical sensor may include an enzymelayer and a polymer membrane.

The polymer membrane may be configured to act as a diffusion barrierthat allows oxygen to go through unhindered but hinders glucosediffusion. The polymer membrane may comprise polyurethane, a mixture ofpolyurethane and silicone, or a mixture of polyurethane and PEG.

The enzyme layer may be under 3500 nm in thickness. Alternatively, theenzyme layer may be less than 1000 nm in thickness. Furtheralternatively, the enzyme layer may be between 200 nm and 800 nm inthickness. Further alternatively, the enzyme layer may be between 600 nmand 800 nm in thickness.

The polymer membrane may be between 200 nm and 10500 nm thick.Alternatively, the polymer membrane may be between 200 nm and 1500 nmthick.

Optionally, the electrochemical sensor may further include abiocompatibility layer. The biocompatibility layer thickness may bebetween 1000 nm and 20000 nm. A biocompatibility layer with a thicknessof less than 1000 nm may be used alternatively.

One or more of the enzyme layer, the polymer membrane and thebiocompatibility layer may be made by spin coating or spray coating.

According to an eleventh aspect of the present disclosure, there isprovided an applicator device for use with the glucose sensor system ofthe first aspect to insert the implantable monolithic integrated circuitunder subject skin, the applicator device comprising: a needle foraccommodating the implantable monolithic integrated circuit and at leasta part of the transcutaneous connector attached to the implantablemonolithic integrated circuit; an actuation mechanism configured to pushthe needle towards the subject skin so as to pierce the subject skin andto insert the implantable monolithic integrated circuit under thesubject skin; wherein the needle comprises a sidewall and a slot in thesidewall, wherein the slot is configured to allow the transcutaneousconnector and the implantable monolithic integrated circuit to escapefrom the needle.

It would be understood that the cross sectional size of the needle issufficient for the needle to receive (surround) the implantablemonolithic integrated circuit and the at least a part of thetranscutaneous connector, and that a width of the slot is sufficient toallow the implantable monolithic integrated circuit and the at least apart of the transcutaneous connector to leave the internal spacesurrounded by the sidewall of the needle.

Advantageously, the applicator device allows the transcutaneousconnector to remain attached to each of the implantable monolithicintegrated circuit and the transmitter during the insertion process. Assuch, a user (e.g., the subject or his/her physician) would not berequired to attach the transcutaneous connector to the transmitter afterinserting the implantable monolithic integrated circuit under thesubject skin. The applicator device may further comprise a retractionmechanism configured to retract the needle from the subject skin.

The applicator device may further comprise a transmitter holderconfigured to support the transmitter of the glucose sensor system(e.g., by friction). The actuation mechanism may be configured to pushthe transmitter holder (hence the transmitter) towards the subject skin.

The applicator device may be operable in a loaded state in which thetransmitter, the implantable monolithic integrated circuit and thetranscutaneous connector are loaded inside the applicator device forinsertion under the subject skin. The transcutaneous connector remainsattached between the implantable monolithic integrated circuit and thetransmitter during the loaded state.

The actuation mechanism may comprise an external cover and a chassisslidably received by the external cover, wherein the needle ismechanically coupled to the external cover such that movement of theexternal cover towards the chassis pushes the needle out of the chassis.

The retraction mechanism may comprise a spring which is configured topush the needle back into the chassis.

The needle may be a hypodermic needle.

The transmitter may comprise a needle guide (e.g., a through hole) whichallows the needle to pass through the transmitter.

According to a twelfth aspect of the present disclosure, there isprovided a method of operating the applicator device of the eleventhaspect, the method comprising: loading the transmitter, the implantablemonolithic integrated circuit and the transcutaneous connector of theglucose sensor system of the first aspect inside the applicator device,wherein the transcutaneous connector remains attached to each of theimplantable monolithic integrated circuit and the transmitter during theloaded state; pressing the applicator device against subject skin topush the needle towards the subject skin so as to pierce the subjectskin and to insert the implantable monolithic integrated circuit underthe subject skin; and retracting the needle by a spring force of aretraction spring of the applicator device.

It would be understood that the transcutaneous connector and theimplantable monolithic integrated circuit escape from the needle via theslot extending through the sidewall of the needle during the retractingstep.

In the present disclosure, the term “implantable” may be usedinterchangeably with “minimally invasive” or “insertable”, and isintended to mean that the respective monolithic integrated circuit (orsensing circuit) is suitable for measuring analytes (such as but notlimited to glucose) in the tissue under subject skin. The subject may bea human being or an animal.

In the present disclosure, the terms “wire”, “transcutaneous connector”,“flexible connector” or “flexible PCB” refer to a structure comprisingconducive traces and insulations to enable taking power and signalto/from the implantable sensing circuit to the transmitter.

In the present disclosure, the terms “applicator”, “inserter” or“applicator device” refer to a device comprising mechanical parts toenable inserting the implantable sensing circuit under the subject skinvia a small needle.

It will be appreciated that features described in the context of oneaspect of the disclosure may be used with another aspect of thedisclosure.

It would further be appreciated that the various numerical rangesdescribed above allow for a degree of variability, for example, ±10%, inthe stated values of the end points of the ranges. For instance, astated limit of 0.5 mm may be any number between 0.5 mm*(1-10%), and 0.5mm*(1+10%). Further, values expressed in a range format should beinterpreted in a flexible manner to include not only the numericalvalues explicitly recited as the end points of the range, but also toinclude all the individual numerical values or sub-ranges encompassedwithin that range as if each numerical value and sub-range is explicitlyrecited.

BRIEF DESCRIPTION OF THE FIGURES

Embodiments of the present invention will now be described, by way ofexample, with reference to the accompanying drawings, in which:

FIG. 1 shows a diagrammatic view of a system according to an embodimentof the disclosure, including an implantable monolithic integrated sensorcircuit 1 connected to a transmitter 2 by wire 3, the transmitter 2connected by wireless connection 6 with a smart reader (e.g.,smartphone, tablet) 4 and an insulin infusion mechanism 9 (e.g.,infusion pump). The same wireless connection or a different wirelessconnection 7 can be used to upload encrypted sensor data in a securecloud 5.

FIG. 2A and FIG. 2B show an embodiment of an implantable monolithicintegrated sensor circuit 1 with multiple on-chip electrodes. FIG. 2Ashows a top view. The monolithic integrated sensor circuit 1 includes acommunication unit 140, a power management unit 120, an integratedsensing element 160, a sensor signal acquisition unit 130, and on-chipconnectors 175 located on an interface unit, silicon substrate 110. Thesensing element 160 comprises working (i.e., WE1, WE2, and WE3),reference (RE), and counter (CE) electrodes. Exemplary dimensions areprovided for the monolithic integrated sensor circuit 1. FIG. 2B shows aside view. The monolithic integrated sensor circuit 1 includes a siliconsubstrate 110 with integrated contact pad 175 attached to flexibleconnector 3, the integrated working electrodes, WE1, WE2, and WE3,representing the integrated sensing element, said working electrodesseparated by insulation walls 118, with the working electrodes coatedwith the functional matrix example Gox-loaded hydrogel 158 and thepolymer coating example PU membrane 168.

FIG. 3 shows a scheme of fabricating the integrated sensor 1, theflexible connector 3, the transmitter 2, and the applicator (inserter) 8and assembling them together in a scalable manner. This fabrication usesa series of process steps commonly used in the semiconductor industryincluding CMOS fabrication, Postprocessing, Wafer thinning, Spincoating, dicing, bonding, enzyme chemistry deposition and spin coating,Polymer (e.g., PU) membrane deposition and spin coating or spray coating(both possible), dicing pf the bonded substrates to separate the bondedsensors, coating of the separated devices with biocompatible material(e.g., PVA) to cover the sidewalls, and assembly of the sensor with thetransmitter and the applicator following sterilization.

FIG. 4 shows a schematic of the components of an implantable monolithicintegrated sensor circuit 1 according to some embodiments of thedisclosure. The implantable monolithic integrated sensor circuit 1 caninclude an integrated sensing element 160, a power management unit 120,a sensor signal acquisition unit 130, and a communication unit whichcomprises a signal receiver unit 1401, a signal transmitter unit 1402,and a MUX/DEMUX unit 1403.

FIG. 5 shows a schematic of the components of an implantable monolithicintegrated sensor circuit 1 according to some embodiments of thedisclosure. The implantable monolithic integrated sensor circuit 1 caninclude an integrated sensing element 160, a power management unit 120,a sensor signal acquisition unit 130, and a communication unit whichcomprises a signal receiver unit 1401, a signal transmitter unit 1402,and a specific example form of MUX/DEMUX in the form of a capacitor.

FIG. 6 describes an embodiment of an element of a sensor signalacquisition unit 130. An ADC is depicted which can convert a sensorsignal (e.g., sensor current) into a digital signal (e.g., bit stream)using a mixed-mode circuit consisting of an amplifier 13213, acomparator 13210, a timing circuit 13212, and a control and logic unit13214. A reset switch RST is used to reset the system before conversionof a new signal from the sensor. A capacitor is used to accumulate thesensor current over a given time period, after which it is converted toa digital signal using the rest of the circuit. The reset switchdischarges this capacitor to restart the process.

FIG. 7 describes an embodiment of working, counter, and referenceelectrodes with corresponding wiring schemes. At the top, an analogfront end (AFE) of a model working electrode includes an amplifier 513which can be used in negative feedback to maintain the potential at theworking electrode, V_(WE), while buffering the sensor current to theADC. The bottom shows an amplifier 512 which can be utilized to formnegative feedback between reference and counter electrodes. Thereference electrode's potential is maintained at V_(RE) through thenegative feedback, while the amplifier provides counter electrodecurrent.

FIG. 8 shows one embodiment of the circuit to generate different WEpotentials for the different on-chip redox reactions. In one embodiment,the V_(WEn)'s are generated using a Digital-to-Analog Converter (DAC).An example of a DAC is shown implemented using a resistor dividernetwork. Utilization of the DAC also allows for programmability of theredox potentials.

FIG. 9 describes an embodiment of sensing circuitry (e.g., theintegrated sensing element 160 and the sensor signal acquisition unit130) using minimal chip area through time multiplexing and sharing ofcounter (CE) and reference electrodes (RE). Multiple separate workingelectrodes (WE₁ . . . WE_(n)) with respective analog front-ends (AFE)can be used for this. Multiple AFEs (e.g., 513 of FIG. 7 ) can be usedas each AFE with the grouping denoted as 514. The multiple workingelectrodes can be managed each by multiplexer connections M₁-M_(n). Theredox potential at each WE is controlled by the corresponding V_(WE)voltage. This enables the design to have different working potentialsfor different working electrodes, thus enabling a wide variety ofanalytes to be detected.

FIG. 10 shows embodiments of different electrode configurations of theworking electrode 12, counter electrode 13, and reference electrode 11that can be manufactured by various fabrication processes according tothe present disclosure.

FIG. 11 shows a temperature sensor circuit. A bandgap circuitrygenerates currents proportional to absolute temperature (PTAT) andcurrent complementary to absolute temperature (CTAT).

FIG. 12 shows an amplification scheme for the temperature sensor circuitcurrent to increase sensitivity.

FIG. 13 shows a diagram of an embodiment of on-chip pads and programmingcircuitry of the monolithic integrated sensor circuit 1 which can beused to store some critical information on the chip itself.

FIG. 14 shows the process of post processing and functionalization forthe integrated sensor 1. The first step shows a side cross-sectionschematic view of a precursor of a monolithic integrated sensor circuit1 as, for example, received from a commercial fab (e.g., TSMC, ON Semi).The schematic shows a substrate 110, a metal-insulator-metal stack 129,a top metal 128, and a top insulator with an optional additionalinsulator/protection polymer layer 118. The second step includes using aprecursor patterned using lithography to create patterned photoresist148. Top first layer metal 128 of a Metal-Insulator-Metal (MIM)structure is Aluminum in some cases.

Underneath the top metal is the further Metal-Insulator-Metal (MIM)structure 129 found in CMOS devices. A silicon substrate under the CMOSstructure is 110. Note that several electrodes can be isolated from eachother by top insulation 118 (Top insulation in CMOS process is often astack of Silicon Nitride layer on top of Silicon oxide layer). Topinsulation can be further augmented by an additional layer (e.g.,Polyimide layer). The third step is a precursor etched to remove the topmetal 128 leaving the area which was protected by photoresist 148 andfurther exposing a second further layer of the Metal-Insulator-Metal(MIM) structure 129. The silicon substrate remains at 110. Electrodesremain isolated from each other by top insulation 118. The fourth stepis a precursor coated with a thin layer of desired material (e.g.,Platinum) 138 using a thin-film coating method, for example thermalevaporation, electron beam deposition, or sputtering. The coating methodis significantly important as it enables fine control on metal surfacemorphology. For example, a sputtering process done at low power (e.g.,200 Watts) and high pressure (e.g., >10 millitorr) can provide a highsurface area Platinum (also known as Platinum Black) like one that isachieved using electroplating in conventional fabrication methods.Moreover, use of different types of plasma conditions (e.g., a mixtureof Argon and Oxygen) plasma in the sputtering chamber enables finecontrol of the surface morphology and coating type. This results in ascalable method to generate black Platinum which has highelectrochemical activity as compared to uniform Platinum deposited usinglow-pressure sputtering or electron beam evaporation. Moreover,controlling plasma conditions (e.g., incorporating Oxygen in the plasma)in the sputtering system can enable formation of desired compounds likePlatinum Oxide (PtOx) which can be a more suitable material forreference electrode. In a different embodiment, a chlorine plasma can beused to convert silver (Ag) into Silver Chloride (AgCl) near the surfacewhich makes a more suitable reference electrode than Ag by itself. ThisPlasma exposure can also be done after the sputtering plasma so surfacereactions are desired instead of reactions during deposition. Asubstrate 110, a metal-insulator-metal stack 129 and a top insulatorwith an optional additional insulator/protection polymer layer 118remains. The top insulation electrode isolation region remains protectedby photoresist 148. The fifth step is a precursor cleaned to remove thephotoresist 148 and excess deposited material and leave depositedmaterial only in desired places (e.g., on integrated sensing elementelectrodes). A substrate 110, a metal-insulator-metal stack 129, a thinlayer of desired material 138 and a top insulator with an optionaladditional insulator/protection polymer layer 118 remains. The sixthstep uses a solenoid and/or piezoelectric controlled actuation-basedspray heads 188 to deposit, in nanoliters, precise amounts of GlucoseOxidase solutions and crosslinking agents (e.g., Glutaraldehyde)solution droplets to make a glucose oxidase-based hydrogel on thesensing electrode. The substrate can be spun after the dropletdeposition to control the thickness of the hydrogel layer moreprecisely. A substrate 110, a metal-insulator-metal stack 129, a thinlayer of desired material 138, and a top insulator with an optionaladditional insulator/protection polymer layer 118 remains. Optionally,the solenoid and/or piezoelectric controlled actuation-based spray heads188 can be used simultaneously with spinning on table to improve theuniformity of the hydrogel. The seventh step shows an embodiment of aprecursor after deposition of the functionalization layer 158 usingsolenoid and/or piezoelectric controlled actuation-based spray heads. Asubstrate 110, a metal-insulator-metal stack 129, a thin layer ofdesired material 138, and a top insulator with an optional additionalinsulator/protection polymer layer 118 remains. The eighth step uses aspray coating (e.g., using fine dispensing heads 186) to make a film(stack of one or more thin films) on a sensor surface to coat thefunctionalization layer, using an appropriate spray head 186 whichshowers microdroplets on an area of the sensing element. A substrate110, a metal-insulator-metal stack 129, a thin layer of desired material138, a functionalization layer 158 and a top insulator with an optionaladditional insulator/protection polymer layer 118 remains. Spin coatingenables a fine control on the thickness of the surface functionalizationlayers. Such layers can be patterned using standard techniques likePhotolithography (e.g., using a sacrificial layer). The ninth step showsan embodiment of a device after deposition of a film 168 on the sensingelement. A substrate 110, a metal-insulator-metal stack 129, a thinlayer of desired material 138, a functionalization layer 158 and a topinsulator with an optional additional insulator/protection polymer layer118 remains. The tenth step shows an instance of how dipping can be usedto apply a small volute of material (e.g., polymer 178) on thesidewalls. A substrate 110, a metal-insulator-metal stack 129, a thinlayer of desired material 138, a functionalization layer 158, a film 168and a top insulator with an optional additional insulator/protectionpolymer layer 118 remains.

FIG. 15 shows a scheme to selectively coat parts of the integratedsensor circuit 1 with multiple layers of chemistry. It shows aphotoresist 181, a sacrificial layer 182, a soft biomaterial (e.g.,enzyme hydrogel) 191, a polymer membrane 192, and a biocompatible layer193.

FIG. 16A, FIG. 16B, and FIG. 16C show schemes of patterning integratedsensor electrodes (of the integrated sensing element 160). FIG. 16Ashows how a sequence of top metal 121 pillar structures can be madeusing semiconductor processes with top insulation 122 covering thespacing in between, by designing the top metal as multiple smallerelectrodes instead of one large electrode. The metal pillar or the topinsulator can then be etched to create open areas between thepillar-like structures. For example, FIG. 16A shows the etching of topmetal to create this spacing. Afterwards, the pillar-like structure canbe coated with suitable metal (e.g., Titanium/Platinum) using thin-filmcoating techniques (e.g., electron beam evaporation, thermalevaporation, sputtering, or atomic layer deposition). It also showsPlatinum coating 125. FIG. 16B and FIG. 16C show two common patternsthat can be used to create high surface area structures. FIG. 16B isbefore etching. FIG. 16C is after etching. After etching of either thetop insulator or top metal, these structures reveal reverse patterns of3D structures (e.g., pillars and valleys or holes and top surface) thatcan then be coated with suitable metal (e.g., Titanium and Platinum). Ina pillar design, the metals are designed as separated electrodes withinsulator filled in the gap areas. In a hole-based pillar design, theelectrodes are designed as metal mesh and the hole is filled by theinsulator.

FIG. 17 shows an example of the integrated sensor circuit 1 with workingand counter electrodes patterned to have pillars structure. It showsthat the pillar structures can have defects which manifest either asbreakage of a set of pillars or coverage of a set of pillars with somedebris from other processes. It also shows the position of interlayervias on the bottom surface (valley area) of the electrode. It also showsthe vias 1291.

FIG. 18 shows an example of the integrated sensor circuit 1 with workingand counter electrodes patterned to have inverted-pillars (holes)structure. It shows that such a structure has much lower defects thanthe pillar structure from FIG. 17 . It also shows the position ofinterlayer vias on the bottom surface (valley area) of the electrode. Italso shows the vias 1291.

FIG. 19 show a comparison between smooth Platinum coated using electronbeam deposition and low-pressure sputter deposition vs. rough platinumdeposited using high-pressure sputter deposition. It shows the smoothplatinum layers of range 30-150 nm (typical 100 nm) vs. rougher platinumlayer of range 30-150 nm (typical 100 nm). A typical smooth layer of 100nm thickness has rms roughness in <3 nm range for 100 nm thickness whilethe rough platinum has rms roughness >3 nm. Moreover, the rough Platinumappears to be porous under an SEM while the smooth Platinum appears morelike a continuous film. The rough surface of the rough platinum gives ithigher surface area on the same geometric area and also makes it moreelectrochemically active. This increases the electrochemical activity ofthe sensor as compared to the smooth Platinum.

FIG. 20 shows a high-level description of an embodiment of a power/dataconnection two-wire scheme 3 between the implantable monolithicintegrated sensor circuit 1 and the transmitter 2.

FIGS. 21A, 21B, 22A, 22B, 23A, 23B, and 24A, 24B, show differentexamples of the implementations of the communication over power. Thesefigures highlight the path in use in each case and show the rest of thecircuit in a lighter tone. For example, FIG. 21A shows signaltransmission from the chip side to the RX circuit (e.g., the amplifierson the transmitter side feeding to an input port of the microcontroller)on the transmitter side via the flexible connector. The TX side of thetransmitter is not active in this state and hence is shown in a lightertone color, and so are the receive circuits on the chip side. Thesignals on the different points of the transmission from the sensorcircuit 1 to the input pin of the transmitter are shown in FIG. 21A. Itshows that the signals starting from the TX side of the chip are perfectdigital signals which decrease in their amplitude as those are coupledwith the power line. Hence, on the RX side of the transmitter, those areamplified and restored back to the original shape (e.g., rail-to-raildigital signal) and are then fed to the digital circuit of thetransmitter for detection and demodulation. FIG. 21B shows a scheme formoving data from the transmitter to the chip.

FIGS. 22A, and 22B show a different coupling-decoupling scheme (i.e.,use different mux/demux circuits for combining and separating the dataand the power signals) in which no capacitor is used on the RX side onthe transmitter but only on the TX side. Also, on the chip side, thecapacitor is used in the TX mode, but no capacitor is used in the RXmode.

FIGS. 23A and 23B show another coupling-decoupling scheme in which nocoupling capacitor is used i.e., neither the transmitter side nor thechip side.

FIGS. 24A and 24B show a scheme in which the power and data lines areseparate. Hence, the TX and RX data on both the transmitter and chipside need to be combined but the power signal flows on a separatededicated wire.

FIG. 25 provides a packaging process scheme of the monolithic integratedsensor circuit 1 onto the connector (or wire) 3 which may be flexible.Step 1 illustrates a sensor circuit 1 positioned adjacent to a flexibleconnector 3 (e.g., a flexible PCB panel). The flexible connector 3 mayinclude conductive traces 302 in a middle layer of the flexible PCBpanel, i.e., under a PCB top layer 303. The sensor circuit 1 can includemultiple sensors (or sensing elements) 160 at some distance from CMOSpads 175. Step 2 shows the bonding of the sensor circuit 1 to theflexible connector 3 at the substrate 301 through die attach, followedby forming conductive traces between the sensor circuit 1 and flexibleconnector 3 via CMOS pads 175 and wire bonds 314. An image of a wirebonded sensor circuit 1 can be seen below.

FIG. 26 shows a continuation of the packaging process scheme. Step 3shows the conductive pads and the conducting interface on the sensorcircuit 1 and the flexible connector array 301 protected for wateringress and mechanical abrasion by covering those with a biocompatibleinsulating material 313. An image of a wire bonded sensor circuit 1 canbe seen coated below. At the bottom of the figure is shown a side viewof the flex PCB shown above. Bonding the implantable monolithicintegrated sensor circuit 1 to the two-wire flexible connector 3 isaccomplished through wire-bonding and insulating the connection usingbiocompatible polymers such as epoxy, parylene-C. 311 shows theconductive pad on Flex PCB 301. 311 makes contact with wire-bond 314which in turn contacts CMOS pad 175. The wire bond can be covered by aninsulating epoxy 313.

FIG. 27 shows a process of wire bond encapsulation followed byfunctionalization in a scalable manner. The first step shows an instanceof how a strip of monolithic integrated circuits 610 (each of which isan example of the monolithic integrated circuit 1) can be attached to aflexible PCB panel 620 (which can be cut to produce multiple instancesof wire or connector 3) using, for example, die attaching followed bywire bonding using wires 630. The second step shows how an insulatingmaterial 640 can be used to hermetically seal the sensor-PCB interface(the pads on the sensors, the wire bonds, and the pads on the PCB). Thethird step shows how one or more soft materials like enzyme hydrogel 190can be coated on the sensors using a thin-film coating process (e.g.,spin coating, spraying). This can be followed by multiple similar stepsto coat a second soft material like polyurethane 189, followed by athird soft material lie Polyvinyl Alcohol (PVA). The fourth step showsan instance of how the individual monolithic integrated circuits 1attached to wire 3 (combination of which is labeled 660) can beseparated by dicing the sensor-PCB assembly together. This can be doneusing different dicing methods like mechanical (e.g., saw) dicing orlaser dicing. This process enables scalability by allowing a strip (canbe extended to have two sensor rows, one on each side of the stripinstead of one) of sensors to be handled together with a larger Flex PCBpanel in a manufacturing process, instead of handling individualmonolithic sensor circuits 1. The materials used (e.g., epoxies, wirebonding wires, functional materials, membranes) are same as those usedfor a single sensor example.

FIG. 28 shows another packaging scheme to attach one monolithicintegrated sensor circuit 1 with the flexible PCB 301 through flip-chipbonding while making the packaged device more planar and more uniform.For this, conductive bumps 1751 are formed on the CMOS pads 175 or onthe flexible PCB conductive traces 302 or both (corresponding padsprovide for stronger and more symmetric bonding). Afterward, the CMOSand Flex connector pads are aligned and pressed together for a certainduration at controlled temperature and pressure, through a suitabledielectric or multilayer stack of the Flexible PCB 301 to minimize theheight difference between the PCB connector and the monolithicintegrated sensor. Typically, the pads on the CMOS sensor circuit 1 arecovered with a metal bump made of soft metal (e.g., Gold or Indium-Tin)to enable reflow and strong conductive connection with the flex pads. Itis possible to add these bumps on the flex PCB conductive traces 302 aswell, although it is not typically required.

FIG. 29 shows a continuation of the packaging scheme. The interfacebetween the flex PCB Pads 302 and the CMOS pads 175 can be covered ininsulating material 313, e.g. using an underfill or a similar process.The zoomed portion of the figure provides a packaging of the implantablemonolithic integrated sensor circuit 1 to PCB trace 302 (e.g., atwo-wire flexible connector). This can be accomplished through bondingthe monolithic integrated sensor circuit 1 to the two-wire flexibleconnector through flip-chip bonding. 302 shows the conductive pad (orconductive trace) on Flex PCB 301. 302 makes contact with a conductivebump 1751. Conductive bump 1751 sits on CMOS pad 175. The interfacebetween the flex PCB 301 and the CMOS pad 175 can be covered ininsulating epoxy 313.

FIG. 30 shows another scheme of connecting the monolithic integratedsensor circuit 1 and the flexible connector 3. In this scheme, thesensor circuit 1 is embedded within the flexible connector (e.g.,flexible PCB) and is electrically connected using small conduciveconnections (e.g., inter-layer vias) 38 which on one side connect to thechip pads 175 and on the other side connect to flexible PCB contact pads39 via conductive traces. This method avoids having to use any extrainsulating materials and use the flexible PCB material (e.g., Kapton,LCP, PEEK, etc.) as the insulating material. Another advantage of thistechnique is that the layer thicknesses can be controlled to result in aplanar device by coating the sensing surface 160 with functionalmaterial 19. Moreover, embedding the sensor circuit 1 within theflexible PCB improves the mechanical robustness of the system. For thisembedding, the sensor circuit 1 is typically thinned down and coatedwith a polymer (e.g., polyimide) layer followed by dicing and embeddingwithin the flexible PCB layers (e.g., using thermal bonding between thebackside Polyimide and the Flexible PCB polyimide). The contact pads atthe sensor circuit 1 are connected with the conductive traces of theflexible connector 3 by via-like metal fills in small holes in theinterlay insulator or by an anisotropic conductive adhesive (ACA).

FIG. 31 shows another scheme of connecting the monolithic integratedsensor circuit 1 and the flexible connector 3 (e.g., flexible PCB). Inthis scheme, the sensor 1 uses through-silicon-via (TSV) technology tohave contact pads 19 under it. In this scheme, a trench is etched fromthe backside (through the substrate 110) of the CMOS sensor circuit 1 tothe conductive trace using a combination of wet and/or dry etchingmethods. Then the sidewall of the silicon is covered with and insulatorand finally with a conductive material that runs from the backside ofthe CMOS sensor circuit to the front side e.g., to chip pad 175). Thesepads on the backside of the CMOS sensor circuit 1 can be attached to theflexible PCB using different methods (e.g., bump bonding, conductingadhesive like Anisotropic conductive adhesive or ACA) depending upon thesize and the application. Next, the conducting interface is covered withan insulating material 313.

FIG. 32 shows another embodiment of the packaging scheme. In this case,the monolithic integrated sensor circuit 1 is placed inside a well inthe flexible connector 3. The gap between the sensor circuit pads 175and the flexible connector pads 302 is filled with an insulatingmaterial 313, followed by forming the connection between the two pads314 (e.g., using wire bonding, or conducting epoxy, etc.) followed bycovering it with a thin layer of insulating material 313.

FIG. 33 shows a scheme of using the IMS integrated sensor circuit 1 onone side of the flexible connector and using the other side of theflexible connector as a secondary sensor 9. The flexible connector 3 onthe other far side uses a connector that can connect with its conductingtraces on both sides, thus enabling connection with both the integratedsensor circuit 1 and the secondary sensor 9. The secondary sensor can beused for validation and control purposes.

FIG. 34 shows a scheme of using an ASIC chip 30 without an on-chipsensor, connected with an off-chip (e.g., on the flex connector 3 withconductive traces 302 that lead to the transmitter 2) sensor to takeadvantage of a hybrid sensor design. The ASIC 30 has contact pads 31 onone side to be connected to a sensor (e.g., Potentiostat contact pads)and contact pads 32 on the other hand which are to be connected to theexternal transmitter. In one example, a 3-electrode sensor consisting offlexible PCB metal layer electrodes 40 terminate into contact pads 42via conductive traces 41. The contact pads are attached to contact pads31 on the CMOS sensor circuit 1 using different methods (e.g., wirebonding, flip-chip bonding) by forming a conductive path 43 between thetwo. The other end of the CMOS chip is also connected to the flex PCBpads 302 via similar methods (e.g., using wire bonding). The conductiveinterfaces are then covered into suitable insulating materials (e.g.,insulating epoxy) 45 and 46.

FIG. 35 shows a high-level schematic of the transmitter 2 according tosome embodiments of the disclosure. The transmitter 2 can include atransceiver 210 that can generate and detect communication signalsflowing over a combined power/data connection through wire 3, ahigh-performance digital microprocessor unit 211 for system control, alow-energy wireless communication chipset 213 (e.g., Bluetooth), anantenna 218 for wireless communication (e.g., Bluetooth antenna), apower management unit 214 connected to a battery 212, and a voltageregulator unit 217 to provide regulated power to the implantablemonolithic integrated sensor circuit 1.

FIG. 36 shows a detailed schematic of an embodiment of transmitter 2which includes a battery 212, a voltage regulator 217, a voltagereference 216, a microprocessor/microcontroller 211, a wirelesscommunication transceiver 213, an antenna 218, wire 3, a DC decouplingcapacitor 242, a switch 228, a switch 229, an Rx unit 210, a Tx unit215, a secondary voltage regulator 231, a voltage reference circuit 232,and resistor 243.

FIGS. 37A and 37B show an implementation of the transmitter circuit 2.FIG. 37A shows a diagram view including sensor 1 with connector 3attached (e.g., by wire bonding, soldering, through a medical grademicro-connector, etc.) with the transmitter 2 which further comprises aprinted circuit board 25 housing the battery 212, regulator 217, andintegrated BLE transceiver and application controller/processorSystem-on-chip 221. The transmitter is housed in a sealed plastic/rubberhousing 22 for environmental protection and easy handling during usage.FIG. 36B shows images further inclusive of connectors 32 and 34, whichcan mate to connect the printed circuit board and wire 3. The wire 3 mayalso be referred to as a “transcutaneous connector” or “flexibleconnector”.

FIG. 38 provides a scheme to package the transmitter electronics insidea sealed casing 22 such that the wire 3 can be connected to thetransmitter PCB 25 using a suitable connector or connector pair 220(e.g., mated pair of connectors 32 and 34) used to connect the flexible(e.g., 3) and rigid (e.g., 25) PCBs. In case of connectors pair, 220represents a mated pair. However, it can also be a single connector(e.g., a ZIF connector) that mates directly with contact pads designedon the flexible connector or flexible PCB 3. This is different thanother transcutaneous CGMs that use soft conductors to mate their sensorswith the transmitter PCB and not standard flexible electronics-basedconnectors used in the IMS design. The figure also shows the other maincomponents of the transmitter electronics including the Analog frontend251, the transmitter System-on-chip or ASIC 250 which contains themicrocontroller along with some peripherals and the communication (e.g.,BLE) chip as well as a power source 212.

FIG. 39 shows an algorithmic scheme to program the processor (e.g., amicrocontroller) in the transmitter (which acts as the brain of thetransmitter) with a firmware to control its operation. It shows that thetransmitter (microcontroller) is programmed to stay in a low power(e.g., deep sleep) state. After the user opens the packaging andperforms a turn-on operation (e.g., either by using an NFC device topair with it) or the transmitter automatically detects the user's intentto use it (e.g., by detecting a change in background conditions likelight via a photodiode), it performs an initialization sequence whichincludes a self-test as well as a scan for the reader via BLE. Once itfinds a matching reader, it connects with it via BLE. Next, it tests ifa good sensor is connected by performing electrical measurements (e.g.,voltage drop, current draw) and by sending a command signal and testingthe response. After that, it starts transmitting the tag signal andstart reading the corresponding data from the sensor chip. Thetransmitter separates the power and data signals (via mux/demux) andsends it to the microcontroller which preprocesses the data (e.g., checkfor proper preamble, proper data coding scheme, packet length, packetduration), checks if it detects any error (via errordetection/correction code), and separates the data from all 4 (3electrochemical and 1 temperature) sensors. It then sends the data tothe reader via BLE.

FIG. 40 shows a scheme for using the multiple on-chip sensors (workingelectrodes) data as well as previous readings and a personalized (foreach patient) patient-sensor model to generate the best outcome (glucosevalue) at a given time. The scheme uses the current data from allsensors (3 electrochemical and 1 temperature). and uses that tocalculate errors among the sensor values. It also compares the readingswith the previous reading to decide if the new values arephysiologically accurate. By comparing the sensor readings amongthemselves and the errors, the system decided if a sensor hasunacceptable level of error. In that case, it discards that sensors anduses rest of the sensors data to generate a weighted average as thecurrent value of the sensor.

FIG. 41 shows the process of insertion of an embodiment of thedisclosure. In step one the monolithic integrated circuit, wire, andtransmitter are placed inside an applicator consisting of an externalplastic body 81, adhesive patch 82, and needle 83. In step two thesensor and injector assembly are pushed towards the skin. This allowsbody analyte to reach the sensor surface and the system to startmonitoring the concentration of one or more analyte in the body.

FIG. 42 shows a different scheme of sensor insertion using a disposableapplicator. The figure shows an exemplary design of such an applicator 8comprising an external cover 81, a triggering guide 82, a smalltriggering arm 83, a needle holder 84, a retracting spring 85, aninsertion needle 86, a transmitter assembly holder 87, and thesensor-transmitter assembly 21 with a needle guide 241. The needle guideallows the needle to pass through the transmitter casing and in turnaround the flexible connector which is attached to the transmitter onone end and to the integrated sensor on the other end. The externalcover 81 also includes two rods 810 extending from the interior of theanterior top towards the bottom posterior open circular aperture. Therods 810 function as guide poles for springs 812 wherein the rods areconfigured to have an exterior diameter smaller than the mean diameterof the respective spring. The rods can extend from bases which can besimilarly shaped to protrusions 88 and can operate as bases for the tipof the coil of springs 812 to sit thereon. The external cover 81 hasthree groves 814 to fit the assembly holder 87 via three features 874.It also has two springs 812 and 813 to provide smooth movement of theassembly. The triggering guide 82 has corresponding holes 821 and 822 tomatch with the springs 812 and 813 of the cover 81. Additionally, it hasthree cut-outs 824 to allow sliding movement of the assembly holder 87via three holders 872 on the assembly holder 87. It also has 4 slots 825to enable alignment of assembly holder 87 with it (triggering guide 82).The needle holder 84 has a holding structure 841 that enables it toslide with the assembly holder 87 within the groves 824 of thetriggering guide 82. The applicator involves internal alignment markers871 and 825, and then mating 3 mm external alignment markers 872 and824.

FIG. 43 shows a step-by-step process of how the applicator 8 is used toinsert the sensor-transmitter assembly 21 on the user body having thesensor connected to the flexible connector under the skin while keepingthe transmitter attached to the skin surface. It shows the first step inwhich the applicator is in a loaded state and is pushed against a bodylocation (e.g., upper arm), after which it inserts the sensor under theskin and retracts the needle (release state). The last part of thefigure shows the arm with the transmitter is on the skin and the sensor1 is under the skin connected via the flexible connector.

FIG. 44 shows the armed (loaded) state of the applicator i.e., the statein which the transmitter-sensor assembly is loaded inside the applicatorfor insertion in the patient. In the loaded stage, the needle 86 is in aposition such that the needle 86 is away from the skin surface and thespring 85 is in the compressed stage. FIG. 44(a) shows a perspectiveview of the applicator 8 during the armed state. FIG. 44(b) shows aperspective view of the applicator 8 during the armed state but with theexternal cover 81 being invisible. FIG. 44(c) shows a cross-sectionalview of the applicator 8 during the armed state.

FIG. 45 shows the next stage of sensor insertion (trigger) in which theapplicator is put on the skin in a desired region (e.g., upper arm) andis pressed against the skin. This pushes the needle holder 84 and thetransmitter holder 872 towards each other. This pushes the needle 86down to pierce the skin. Once and the triggering arm 83 reaches themaximum depth, this defines the maximum depth of the needle under skin.

FIG. 46 shows the next stage of sensor insertion (release) in which theneedle is retracted back to the applicator assembly by the spring forceof retraction spring (85 in FIG. 42 ) after the triggering arm 83reaches the maximum depth in the triggering guide in the transmitterassembly holder. This needle retraction mechanism ensures the needledoesn't cause any accidental injury afterwards. At this step, thebiocompatible adhesive attached to the transmitter-sensor assembly 21keep the transmitter attached to the skin which ensures that the sensorremains under the skin at the desired depth and is connected to thetransmitter via the flexible connector 3. The needle 86 is retractedinside the applicator at this stage.

FIG. 47 shows an example of the sensor-transmitter assembly 21 on theskin with the flexible connector 3 passing through the skin(transcutaneous) after the applicator performs its operation and isremoved from the skin. It also shows the cut-out of the needle guide 241which ensures the needle passes around the sensor such that the sensorsits within the needle body through the slot in the needle.

FIG. 48 shows a side view of the trigger state of the applicator and howthe triggering arm 83 moves past the lowest point (maximum depth) in thetriggering guide when the user pushes the applicator down on the skin sothat the needle 86 can pierce the skin. It also shows the optionalcurves 89 on the top cover that are designed to make it easier to holdthe applicator. As the triggering arm moves just past the maximum depthin the triggering guide, it is pulled up by the spring force in the nextstage (release stage).

FIG. 49 shows a close-up view of how the sensor 1, the connector 3, andthe transmitter (with a PCB 25 inside it) are assembled such that thesensor-connector assembly passes through the cut needle duringapplicator operation.

FIG. 50A shows a view of an actual applicator built using the designpresented here and loaded with a sensor-transmitter assembly with thesensor-connector part of the assembly passing through the applicatorneedle. FIG. 50B shows an image of an actual applied sensor assembly.

FIG. 51A, FIG. 51B, and FIG. 51C show peroxide response of a sensorduring repeated testing. FIG. 51A shows peroxide concentration versussensor current for a sensor with a lower current range (designed in thepotentiostat) while FIG. 51B shows peroxide concentration versus sensorcurrent for a sensor with higher range. FIG. 51C shows example peroxideresponses of three different sensors on the same chip. The smallvariations can be decreased using more repeated processing and cleaningof the sensor surface.

FIGS. 52A, 52B, and 52C show glucose concentration versus sensorcurrent. FIG. 52A shows a chart for sensor one chip. FIG. 52B shows achart for sensor two. FIG. 52C shows a chart for sensor three.

FIG. 53A, FIG. 53B, and FIG. 53C show an in vitro glucose titration foreach of three sensors on another sensor chip. FIG. 53A shows theresponse curve of sensor one. FIG. 53B shows the response curve ofsensor two. FIG. 53C shows the response curve of sensor three.

FIG. 54A and FIG. 54B show in vitro glucose response for each of threesensors on two separate dies, along with their median and averagereadings. FIG. 54A shows data from all 3 sensors on one integratedsensor chip while 54B shows the data from all 3 sensors on a secondintegrated sensor: together with average and median of all 3 sensorsfrom respective dies in both cases.

FIG. 55 shows how the reading from all 3 sensors can be used together togenerate a more accurate overall result than any single sensorindividually using a statistical combination of the sensors” data (e.g.,median in this figure).

FIG. 56 shows the effect of hydrophilic coating (PVA) on sensorresponse. It shows that the PVA coating can be done both with andwithout the polymer membrane (e.g., PU) and it doesn't impact theglucose sensing properties of the sensor negatively.

FIG. 57A and FIG. 57B show data from the integrated temperature sensor.FIG. 57A shows the simulation results for the sensor response. FIG. 57Bshows one reason why the temperature sensor data is required as theenzyme activity is proportional to temperature, leading to change incurrent if temperature changes.

FIG. 58A and FIG. 58B show the testing results for the integratedtemperature sensor. FIG. 58A shows a 7-day test of the sensor in asaline solution to track changes in the environment temperature (dailytemperature cycle). FIG. 58B shows how the sensor can track the changein environment temperature after insertion in a person vs. the roomtemperature before insertion, as well as how it can detect a change inthe tissue temperature created by blowing cold air on the skin.

FIG. 59A and FIG. 59B show human data of the temperature sensor. FIG.59A and FIG. 59B show the temperature sensor can detect fall insubcutaneous tissue temperature due to decrease in glucose and candetect increase in subcutaneous temperature as glucose excursion drivesthis temperature up. This is consistent among two different sensors wornby the same person.

FIG. 60A, FIG. 60B, and FIG. 60C show the IMS glucose sensor response ina First-in-Human (FIH) study. FIG. 60A and FIG. 60B show that twodifferent IMS sensors can trace glucose excursions vs. a contour meter(primary references) as well as commercial CGM references (Dexcom G6,Abbott Libre 2). The Clarke-error grid of FIG. 60C shows the IMS sensorfollowing the contour reference quite accurately.

FIG. 61A and FIG. 61B show the response of the FIH sensor before andafter the FIH study. FIG. 61A shows the sensor was responsive to glucosein the target range for a prediabetic individual (50-200 mg/dl). FIG.61B shows that the sensor was still responsive without a significantloss of its performance after the FIH study (for a 3-day wear period).

FIG. 62A, FIG. 62B, and FIG. 62C show the human testing data from a3-day wear period. FIG. 62A shows day 1, FIG. 62B shows day 2, and FIG.62C shows day 3.

FIG. 63 shows the effect of a simple multi-sensor data fusion scheme inimproving the overall outcome of the sensor. It shows that the datafusion scheme results in a more accurate output, even in cases wheresome of the on-chip sensors have some error.

FIG. 64 shows the improvements in rate responsiveness for the IMS deviceas compared to the commercial references, due to the thin sensorcoatings and higher accuracy (requiring lesser averaging)

DETAILED DESCRIPTION

The present disclosure is directed to an implantable monolithicintegrated sensor circuit and system that can be used in a variety ofin-vivo applications providing continuous measurement of one or moretypes of health or biological markers (e.g., metabolites). Glucose is anexample analyte discussed herein. As a person having ordinary skill inthe art will appreciate, the described devices, systems and methods canbe more generally applied to other analytes and analyte combinations.Additionally, some ex-vivo uses are easily envisioned.

System Design

An overview of an embodiment of the system of the present disclosure canbe seen in FIG. 1 . The system can include a monolithic integratedsensor circuit 1 that is connected to an external transmitter 2 througha connector 3 made of conductive material which may be coated with abiocompatible polymer. The connector 3 may be a flexible connector. Thetransmitter 2 provides electrical power to the monolithic integratedcircuit 1 via the flexible connector 3. It also communicates to themonolithic integrated circuit 1 via the flexible connector 3. Thecommunication can be bi-directional or unidirectional. Transmitter 2 canbe wirelessly linked to a smart reader (e.g., smartphone) 4, smartinsulin pump/pen 9, and/or a secure database 5 using a wirelesscommunication technology such as Bluetooth, Zigbee, or WiFi 6, 7.Alternatively, the smart reader 4 can act as a bridge between thetransmitter and the secure database 5. The details of the importantparts of the system are provided next.

Monolithic Integrated Sensor Circuit

In accordance with some embodiments of the disclosure, the monolithicintegrated sensing circuit is, for example, an integrated circuit chipfabricated using CMOS fabrication technologies known to the personskilled in the art.

An important element of the monolithic integrated circuit is the smallsize of the chip. The monolithic integrated sensor circuit can includemany interconnected functional modules or subsystems and can be in arange from 30 microns to 600 microns in thickness (e.g., 50 microns to150 microns), 500 microns to 10,000 microns in length (e.g., 1500microns to 3000 microns) and in a range from 100 microns to 4,000microns in width (e.g., 400 microns to 1000 microns).

The small size of the monolithic integrated sensor circuit along withshaping can minimize scar tissue formation in the body to a point whereit only helps in keeping the system position stable but does notsignificantly isolate the implantable monolithic integrated sensorcircuit from accessing body fluids. This allows real-time measurement ofimportant analytes (e.g., metabolic glucose level) for criticalapplications requiring instant changes to be reported as soon aspossible (e.g., for hypoglycemic diabetic patients).

Designing the monolithic integrated circuit in accordance with thespecific implantation/insertion site (tissue orientation etc.) can helpin reducing post-implantation complexities. For example, forimplantation/insertion in biological tissues, the sensing platform canbe shaped to minimize sharp edges to minimize tissue damage and henceimmune system response. The monolithic sensor circuit can be shaped tobe longer in one dimension and much smaller in other dimensions toinject or insert the monolithic circuit using very small needles. Thisalso allows the monolithic circuit to fit within the subcutaneous orsubdermal space more easily. Minimizing the device thickness and coatingit with a biocompatible soft material can also make it more flexible andreduce tissue damage.

A precisely controlled minimization of solid-state sensor size alsoreduces detection noise levels and can increase the Signal-to-Noiseratio (SNR), thus improving the sensitivity of the sensor. Thisminiaturization and accompanying SNR improvement is not possible withoutthe added on-chip circuitry which can read the low current without muchadded noise of a longer-distance transmission. Furthermore, a compactintegrated design minimizes contact resistance and capacitance betweenthe sensor and the electronics, further enhancing sensitivity byimproving the SNR of the sensor. Moreover, the decrease in electrodesize reduces its capacitance which further reduces non-faradaic(charging) currents, thus improving SNR and decreasing the time it takesfor the sensor to stabilize. The minimum SNR for a reliable detection istypically considered to be 3 (Signal to Noise Ratio: unitless). Inglucose oxidase functionalized embodiments, when in patient, theintegrated sensor circuit of the present disclosure is capable of SNRranges of 5-30, or more preferably 10-20. In post-processed form, whenin peroxide, the integrated sensor circuit of the present disclosure iscapable of SNR ranges of 5-100, or preferably 60-100, or more preferably70-80. (Please see, Donald M. Morgan and Stephen G. Weber, Noise andSignal-to-Noise Ratio in Electrochemical Detectors, Anal. Chem. 1984,56, 13, 2560-2567, herein incorporated by reference in its entirety).

Sensor fabrication can start with submitting the chip design files to asemiconductor manufacturer (e.g., TSMC (Taiwan), ON Semiconductor(Phoenix, AZ)). The standard semiconductor fabrication processes cangenerate standard wafers of certain sizes (e.g., 12-inch diameterwafers). To reduce the dimension of the device, the original thickness(e.g., 750 μm) of the semiconductor wafer can be thinned down (e.g., to50-250 μm) through mechanical grinding, chemical and/or mechanicalpolishing or chemical etching (e.g., Xenon Difluoride (XeF₂) etchingfrom backside). This step can be done before or after surfacefunctionalization and membrane chemistry deposition. Once thinned, thesilicon becomes more flexible and can improve the integration of thesensor device within the surrounding tissue and reduce foreign bodyresponse. Thinning and/or grinding can be performed by a thinning andgrinding facility (e.g., Advanced International Technologies,Quick-Pak). Some common exemplary CMOS process nodes which can be usedfor fabrication of the monolithic integrated sensor are TSMC 180 nm, 65nm, 55 nm, 250 nm, 90 nm.

Different types of dicing methods (saw, laser, stealth, etc.) along withsome polishing methods can be used to realize any desirable shape (e.g.,circular, rectangular, oval). Laser cutting can be used to form roundededges on the final diced device and help reduce potential implantationinjury and subsequent foreign body response. Laser dicing can beaccompanied by appropriate environmental condition (e.g., oxygen flow)to create a thin layer of thermal oxide on sidewalls during dicing.Steam can also be used to generate a wet oxide on sensor sidewalls.Sidewall polishing after dicing can also be used to reduce and removesharp edges. Further, coating with biocompatible membranes can also beused to minimize any sharp edges.

FIG. 3 shows an embodiment of a complete wafer with multiple rows of themonolithic integrated sensor circuits fabricated using a CMOS process.FIG. 3 and FIG. 27 show an instance of how a row of monolithicintegrated circuits 610 can be separated from the wafer using dicing.Note that, individual or rows of monolithic integrated sensor circuitcan be diced before post processing or as a final step after processing.FIG. 3 and FIG. 27 show an instance of how finished individualmonolithic integrated circuits 660 can be separated by dicing thesensor-PCB assembly together.

An overview of an embodiment of the monolithic integrated sensor circuit1 of the present disclosure can be seen in FIG. 2A and FIG. 2B. In someembodiments of the disclosure, the monolithic integrated sensor circuit1 can include an integrated sensing element 160. The integrated sensingelement 160 may also be referred to as an “electrochemical sensingelement”. The monolithic integrated circuit can further contain a powermanagement unit 120, a sensor signal acquisition unit 130, acommunication unit 140, and an optional interface unit 175. Theinterface unit may provide further circuitry to better communicate withother components of the system, e.g., the transmitter. The interfaceunit is most often wired in the present instance. In its simplestembodiment, the interface unit may consist of a couple of contact padsthat provide a convenient interface for the wired connector which linksthe transmitter. The monolithic integrated sensor circuit includeson-chip connectors 175 in the interface unit. Exemplary dimensions arealso included at 35 μm and 0.15 mm.

A more specific design view of an embodiment can be seen in FIG. 4 . Themonolithic integrated circuit again includes an integrated sensingelement 160, sensor signal acquisition unit 130, power management unit120, and a communication unit 140. The communication unit 140 furthercomprises a receiver subsystem (RX) 1401, a transmitter subsystem (TX)1402, and a MUX/DEMUX network 1403 to separate the communication signalfor the power signal. FIG. 4 shows an interface unit comprising twoelectrical pads. One pad functions as ground from the transmitter. Theother pad functions as data over power connected to the MUX/DEMUXnetwork 1403.

FIG. 5 contains a specific example wherein data is separated from power.The interface unit comprises three pads. A pad is connected to groundinto the power management unit 120. A pad is connected to power into thepower management unit 120. A pad functions as data connected to theMUX/DEMUX network.

As shown in FIG. 4 and FIG. 5 , the receiver subsystem 1401 can composeof a signal detector, a demodulator/detector, and control logic. In thereceiver subsystem 1401, the signal detector (e.g., envelope detector)can be used to extract the data transmitted from the externaltransmitter through, for example, a two-wire flexible connector.

A data encoding scheme (such as pulse width modulation or coding, pulseinterval modulation or coding, pulse code modulation or coding, pulsecount modulation or coding, Manchester Coding etc.) is used to send datafrom the transmitter to the sensor. Pulse code modulation being apreferred embodiment. Therefore, a data demodulator/decoder is utilizedin the sensor receiver to decode the received signal which can includean activation tag for the implant as well as the sensor currentmeasurement range. The control logic can perform signal conditioning andinterpretation of the received data from the external transmitter 2.

As shown in FIG. 4 or FIG. 5 , the transmission subsystem 1402 caninclude a preamble/encoder, serializer, and modulator. The TX unit cantake sensor data from the sensor signal acquisition unit (e.g., ADC),encode it using a specific data encoding scheme for minimizingcommunication error (e.g., Pulse count coding), add predefined sequences(e.g., preambles, pilot sequences) and transmit the encoded data to theexternal transmitter using data communication-over-power scheme.

In the transmission subsystem 1402, the preamble/encoder can combine thesensor data into one or more packets that can be sent to the externaltransmitter. The packetized data can include the sensors (e.g.,electrochemical, temperature) measured data and power level indicator.For example, the preamble/encoder can, in embodiments, combine all thedata elements into a single data packet and add a preamble sequence atthe beginning of the data packet for ease of detection by the externaltransmitter 2. The serializer can serialize data packets received fromthe preamble/encoder. An error detecting and/or correcting sequence(e.g., cyclic redundancy check or CRC, hamming code) can be added to thepackets for immunity to communication and detection noise. The modulatorcan take the data in digital form and change into waveform for sendingto the transmitter 2 over the wire 3 of the implantable sensor circuit.

As shown in FIG. 4 or FIG. 5 , the power management unit 120 can includea regulator (e.g., low V-LDO regulator), reference generator, powerdetector, temperature sensor, and control logic. In some embodiments thepower management may also include a voltage limiter. The regulator canbe a low-dropout regulator that regulates the incoming supply signalfrom the transmitter through a two-wire connector (which is an exampleof the wire 3) into a clean DC voltage (without ripples).

The reference generator can generate the reference voltages and currentsused by the ADC, potentiostats, and the oscillator of the sensor signalacquisition unit 130. The reference generator can provide high powersupply rejection to eliminate sensitivity to supply ripple. Althoughbelow the sensor signal acquisition unit 130 is described as optionallycontaining the digital to analog converter of FIG. 8 to allow thegenerated reference voltages to be programmable, in alternativeembodiments, the reference generator of the power management unit 120can comprise the digital to analog converted which allows the generatedreference voltages to be programmable.

A temperature sensor may be included in embodiments of the powermanagement unit of FIG. 4 or FIG. 5 . An implementation of a temperaturesensor can be seen in FIG. 11 and FIG. 12 . FIG. 11 shows a temperaturesensor circuit. A bandgap circuitry generates currents proportional toabsolute temperature (PTAT) and current complementary to absolutetemperature (CTAT). As shown in FIG. 11 : k—Boltzmann, q—charge e,T—temperature, ln—natural log, n—ratio of the 2d area between the twobipolar transistors. FIG. 12 shows an amplification scheme for thetemperature sensor circuit current to increase sensitivity. To increasethe dynamic range of the temperature sensor, the offset of thetemperature dependent current is removed so that at the temperaturerange of interest the current varies between a value close to zero andthe full range of the ADC. Multiple current amplification stages bringthe range of the current variation close to the ADC dynamic range. Thebottom corner of FIG. 12 shows a mathematical representation of theamplification method implemented.

Another method to implement the temperature sensor is via a resistancetemperature detector (RTD). In this case, a thin filament like electrodecan be made on to of the CMOS device (e.g., on the top layer) using amaterial with good temperature sensitive resistance. The circuitunderneath can read the resistance and hence any change in temperature.An example of the RTD is a Platinum based RTD. Since the IMS systemenables fabrication of multiple electrodes on the CMOS substrate,fabrication of such a temperature probe is simple and can be donetogether with the fabrication of the electrochemical sensors. The RTDprobe can be coated with a thin insulator layer afterwards (e.g., thinSilicon Nitride layer).

Various implementations of the reference generator of FIG. 4 or FIG. 5may include a diode to provide a reference current. At differenttemperatures the performance of the diode may vary. Thus, a temperaturesensor provides a means to measure the temperature and correct for theperformance of the diode within the power management unit. Thesetemperature measurements can be used outside of the power managementunit. For instance, the temperature measurements can be provided to theuser via the transmitter to calibrate the signal obtained from thesensing element 160. The temperature measurements themselves are alsouseful indications of the glucose concentration within a patient's body.This is because a higher glucose concentration tends to trap heat withinthe patient's body, causing the patient's temperature to rise.

The power detector of FIG. 4 or FIG. 5 can be used to measure theincoming power signal level and determine whether the implant isunderpowered, properly powered, or overpowered, and report the powerdata to the external transmitter as part of packet data.

A control logic can be implemented within the power management toexecute the tasks of the regulator, reference generator, power detector,temperature sensor, and voltage limiter, for instance a processor. Thecontrol logic can, in various embodiments, execute tasks for the sensorsignal acquisition unit or communication unit.

In embodiments, the power management unit may include a voltage limiter.A voltage limiter can massage power to be more usable by the monolithicintegrated circuit. For instance, the voltage limiter can protect thesystem from over-voltage by using different methods including by sinkingmore current and hence reducing the supply voltage from the transmitter.In accordance with some embodiments of the invention, an implantablemonolithic integrated sensor circuit 1 can be powered through a powermanagement using the two-wire flexible connector and the datatransmission can use a low-power wireless communication scheme (e.g.,Bluetooth Low Energy, ANT, Zigbee) as implemented through the optionalinterface unit.

FIGS. 21, 22, and 23 illustrate examples of different implementations ofthe MUX/DEMUX (Couple-Decouple) network as well as power and ground bothat the sensor side as well as the transmitter side.

In FIG. 21 a capacitive decoupling network is utilized at thetransmitter and the sensor to decouple the data from the power signal.At the transmitter side an amplifier and a comparator are following thenetwork to amplify and resolve the data sent from the sensor.Communication is performed in a time multiplexed fashion, meaning whensignal is being sent from the transmitter, the sensor is in the receivemode, once the transmitter sends out its data, it turns into receivemode to collect sensor data. This is shown in FIG. 21 in color codeshowing one direction of communication in black and the other in gray.FIG. 21A shows the scheme for sensing data from the chip (i.e., thesensor circuit 1) to the transmitter (i.e., the transmitter 2) and FIG.21B shows the reverse scheme, i.e., from the transmitter to the chip.

FIG. 22 illustrates another variant of capacitive decoupling of the dataand power at the transmitter and the sensor side. In this case, thecapacitor is positioned differently with respect to the transmit andreceive circuitry and the input is fed to the transmitter circuitdifferently as compared to the scheme in FIG. 21 . FIG. 22A shows thescheme for sensing data from the chip to the transmitter and FIG. 22Bshows the reverse scheme, i.e., from the transmitter to the chip. Theanalysis and modeling results show that this scheme can also be sued forcommunication.

FIG. 23 illustrates a resistive decoupling network instead of thecapacitive schemes shown in FIG. 21 and FIG. 22 . In this case, aresistor along with a switch is used to transmit data commands from thetransmitter to the chip and vice versa. FIG. 23A shows the scheme forsensing data from the chip to the transmitter and FIG. 23B shows thereverse scheme, i.e., from the transmitter to the chip. The modeling andsimulation results show that the data can be reliably coupled anddecoupled with this scheme.

FIG. 24 illustrates a three-wire system in which two wires are utilizedto carry power and ground signals and the third wire is used for datacommunication to and from the sensor. Data communication isbidirectional in a time multiplexed fashion. When transmitter is sensingdata, the sensor operates in receive mode. Once data transmission isdone by the transmitter, it turns to receive mode to collect data fromthe sensor. FIG. 24B shows the scheme for sensing data from the chip tothe transmitter and FIG. 24A shows the reverse, i.e., from thetransmitter to the chip.

As shown in FIG. 4 or FIG. 5 , the sensor signal acquisition unit 130can include an oscillator, potentiostat(s), an analog to digitalconverter (ADC), and a multiplexer (MUX) (as well as an optional digitalto analog converter DAC to externally control electrode redoxpotential). The oscillator can be used to provide an accurate and cleanreference clock for the implant that is used by the communication unit,power management unit, and the sensor signal acquisition unit.

The potentiostat is connected to on-chip electrodes. The potentiostatmaintains a fixed defined voltage between working electrode(s) and areference electrode while providing current through a counter electrode.Each working electrode can have a dedicated potentiostat to maintain theappropriate voltage between the working and reference electrode andminimize crosstalk. However, a shared potentiostat is also workable. Ina preferable embodiment, regarding the potentiostat, multiple workingelectrodes share common reference and counter electrodes. In suchinstance, an op-amp (e.g., 512 in FIG. 7 ) can control the referenceelectrode voltage while providing source/sink capability at the counterelectrode and another op-amp (e.g., 513 in FIG. 7 ) can be utilized tocontrol the potential for each working electrode where the amplifiersets the working electrode voltage through establishing negativeresistive feedback and converting the sensor redox current into avoltage for subsequent processing.

An embodiment of wiring schemes of the sensor signal acquisition unit(in particular the potentiostat) connected to working, counter, andreference electrodes can be seen in FIG. 7 . At the top, an analog frontend (AFE) of a model working electrode includes an amplifier 513 whichcan be used in negative feedback to maintain the potential at theworking electrode, V_(WE), while buffering the sensor current I_(SENSOR)to the ADC. The sensor current I_(SENSOR) is inversely proportional tothe resistance of the working electrode which varies with the glucoseconcentration in the blood. The bottom shows an amplifier 512 which canbe utilized to form negative feedback between reference and counterelectrodes. The reference electrode's potential is maintained at V_(RE)through the negative feedback, while the amplifier provides counterelectrode current. The AFE and the amplifier 512 may be collectivelyreferred to as the potentiostat. Therefore, the potentiostat senses thechanges in the resistance of the working electrode, by outputting acurrent which is inversely proportional to the varying resistance of theelectrode. This method allows for independently controlling thepotential difference between working and reference electrodes in amulti-analyte sensor where there are several working electrodes forsensing different analyte.

The potentiostat is continuously powered by the battery 212 of thetransmitter 2 (shown in FIG. 35 ) via the connector 3. By continuouslypowering the potentiostat, the potentiostat is able to maintain thepotential of the working electrode continuously. After installation ofthe battery 212 into the transmitter 2, it may be required to calibratethe system of FIG. 1 once. However, after the initial calibration, thereis no need to further calibrate the system before the battery 212 isdrained up. Therefore, the system of FIG. 1 is able to continuouslymeasure the glucose level of a patient during the lifetime of thebattery 212. The measurement frequency may be from once every tens ofseconds to once every few minutes, and may be adjusted at thetransmitter side (e.g., by a user using the smartphone 4 wirelesslyconnected to the transmitter 2). Note that it is still possible tosupplement the initial calibration with second, third, or furthercalibration points to enhance accuracy or in the event the user believesdeviation has occurred.

The potentiostat is placed at a depth of 2-3 mm beneath the patientskin. The 2-3 mm depth has been found to allow the potentiostat togenerate a superior signal indicative of glucose concentration. The 2-3mm depth also significantly shortens the communication distance betweenthe potentiostat and the transmitter 2. It would be understood that thepotentiostat may be placed at a depth of 1 to 5 mm beneath the patientskin or a depth of 1 to 10 mm beneath the patient skin.

With further reference to FIGS. 4 and 5 , the potentiostat can beconnected within 9, 5, or 2 millimeters of the entirety of at leastthree working electrodes WE1, WE2, and WE3. This means that, forexample, when WE1 is arranged closest to the potentiostat within theworking electrodes WE1 to WE3 and WE3 is arranged furthest from thepotentiostat, the distance between the potentiostat and all the area ofthe furthest electrode WE3 is within 9, 5, or 2 millimeters.Alternatively, the potentiostat is connected within 0.5 millimeter or400 microns to the entirety of working electrodes WE1, WE2, and WE3. Theshort connection distance between the working electrodes and thepotentiostat allows for minimal electronic signal travel, which in turnimproves the signal quality and reduces the power consumption of thesystem of FIG. 1 . It should be pointed out that other continuousglucose monitoring systems have substantially larger distances betweenthe potentiostat and the working electrodes. For example, the Dexcom G6contains a wire that is at least 15 mm long, and thus, the distancebetween the entirety of the working electrode and the potentiostat ofthis device is far greater than those proposed in the instantdisclosure.

A dual slope ADC can be used to directly convert the sensing elementcurrent coming from the potentiostat into the digital domain. The ADCcan, for example, include an 8-12 bit ADC that converts the potentiostatcurrent into digital data values. Since in-vivo glucose readings don'tchange rapidly, each electrode pairing can be sampled every millisecondto once every 100 milliseconds, but preferably once every 10milliseconds (100 samples/second). An ADC with high sampling rate (1Ksamples/s) can be utilized to sequentially digitize the signal frommultiple electrodes as well as the temperature sensor.

An embodiment of an ADC of a sensor signal acquisition unit can be seenin FIG. 6 . An ADC is depicted which can convert a sensor signal (e.g.,sensor current) into a digital signal (e.g., bit stream) using amixed-mode circuit consisting of an amplifier 13213, a comparator 13210,a timing circuit 13212, and a control and logic unit 13214. A resetswitch RST is used to reset the system before conversion of a new signalfrom the sensor. A capacitor is used to accumulate the sensor currentover a given time, after which it is converted to a digital signal usingthe rest of the circuit. The reset switch discharges this capacitor torestart the process.

A multiplexer (MUX) in the sensor signal acquisition unit can compareand compile data from multiple working electrodes on the monolithicintegrated circuit. In accordance with some embodiments, to supportmulti-analyte sensing without an excessive increase in powerconsumption, resource sharing can be enabled across the sensor signalacquisition unit by the multiplexer. In some embodiments of thedisclosure, each individual working electrode can be controlled by adedicated potentiostat while an analog-to-digital converter can beshared among all potentiostats through time division multiplexing inwhich the digitization period is divided among some or all the workingelectrode-potentiostat pairs. During each time slot, the output of oneworking electrode-potentiostat pair is digitized. In accordance withsome embodiments of the disclosure, the sampling rate can be set to arate that is well above the rate at which relevant physiological bodychanges occur to avoid sensed signal loss. Normally, the ADC can operateat a much faster rate than that of the physiological signals, hence suchmultiplexing doesn't create any loss of needed data.

An embodiment of multiplexing sensing circuitry can be seen in FIG. 9 .Multiplexing and sharing of counter (CE) and reference electrodes (RE)is shown. Multiple separate working electrodes (WE₁ . . . WE_(n)) withrespective analog front-ends (AFE) can be used for this. Multiple AFEs513 (e.g., based upon op-amps 513) of FIG. 7 can be used as each AFEwith the grouping denoted as 514. The multiple working electrodes can bemanaged each by multiplexer connections M₁-M_(n) to feed theircorresponding signals (data) to the ADC in a time-multiplexed manner.

The redox potential at each WE can be controlled by the correspondingV_(WE) voltage (which may also be referred to as a reference voltage).This enables the design to have different working potentials fordifferent working electrodes, thus enabling a wide variety of analytesto be detected. In particular, each of these op-amps 513 is connected tothe desired redox potential (e.g., V_(WE1), V_(WE2), V_(WE3), and so on)for that WE. These voltages can be either internally generated as afixed value for each application (e.g., using on-chip reference voltagegenerators), or can be generated using a Digital to Analog Converter(DAC) so that they can be programmed (FIG. 8 ) by the user from a givenset of values to allow for programmability in choosing what analyte tomeasure. The user input of given set of values (also referred to as“command”) may be transmitted from the transmitter 2 to the implantablemonolithic integrated sensor circuit 1 via the connector 3.

An example of the DAC circuitry can be seen in FIG. 8 including aresistor ladder comprising 2^(m) resistors connected between the voltageregulator output and the ground to form an m-bit DAC capable ofgenerating 2^(m) different voltages. These voltages are connected toeach working electrode through a set of switches controlled by D_(ij),where i represents WE_(i) and j is between 1 and 2^(m) for selecting thedesired voltage level. For each working electrode there are 2^(m)dedicated switches.

As shown in FIG. 4 and FIG. 5 , in some embodiments of the disclosure,the implantable monolithic circuit can include an electrochemicalintegrated sensing element that comprises multiple working electrodes(e.g., a detection reaction can occur at this electrode), a singlecounter electrode (e.g., can be used to balance the current generated byworking electrode) and a single reference electrode (e.g., to provide astable voltage reference signal inside the body). The electrochemicalintegrated sensing element transduces changes in analyte concentrationinto changes of electrical current. In an embodiment the integratedsensing element includes three working electrodes, one counterelectrode, and one reference electrode. The sharing of a single counterand a single reference electrode does not create crosstalk as thecurrent in each potentiostat can be determined by the correspondingworking electrode. However, in a separate optional embodiment onecounter and one reference electrode can be used corresponding to eachsingle working electrode. In an example, the desired redox potential maybe between −1.5V to 1.5V. In the event that the resistor ladder of FIG.8 is connected between a positive voltage and a negative voltage, thedesired redox potential may be between −1.5V to 1.5V.

The monolithic integrated sensor circuit 1 can be programmable. Anexemplary programming circuitry of the monolithic integrated sensorcircuit 1 is shown in FIG. 13 . In FIG. 13 , the sensor front-end refersto the potentiostats connected to the integrated sensing element 160 asdescribed above. The process of narrowing down the parameters associatedwith some chip functions (e.g., temperature sensor, electrochemicalreadout circuitry, and the ADC) is also called trimming. As shown inFIG. 13 , an embedded memory is electrically connected to the sensorfront-ends, the ADC and the temperature sensor so as to provide thetrimming parameters to those blocks. The embedded memory is an on-chipmemory and can be programmed via an on-chip memory programmer which isconnected to an external memory controller (not shown in FIG. 13 ) byon-chip pads (e.g., 175). During the electrical testing and trimming,the on-chip pads interface with the external memory controller. Theinformation from the external memory controller can be written into theembedded memory using the memory programmer, which generates highvoltage signals required to program the embedded memory. In some othercases, the memory programmer may be omitted and direct electricalconnections may be established for the external memory controller toprovide correct signals to program the embedded memory. For example, thereference voltages V_(Wei) (i=0, 1, 2 . . . n) of the working electrodes(described with reference to FIG. 7 , FIG. 8 , and FIG. 9 ) may beprogrammed during the above trimming process.

The embedded memory mostly only requires to be programmed once. Hence,simpler one-time programmable (OTP) memories can be used (e.g, opticalROM, EPROM) as the embedded memory. However, in some cases if moreflexibility is desired, a reprogrammable memory (e.g., EEPROM, eFlash)can also be implemented as the embedded memory. The exact type of memoryused depends upon both the application and the CMOS process. The CMOSfoundries (e.g., TSMC) offer a range of memory options for differentCMOS processes that can be chosen appropriately.

Alternatively, optical beams (e.g., lasers) can be used to program theembedded memory without requiring the use of the on-chip pads.

Individual electrodes can be defined by openings in a top passivationlayer (e.g., 118 in FIG. 14 ) of the CMOS fabrication process. Anintegrated design wherein electrodes are defined by the top mostAluminum metal layer in the CMOS process can be used. However, sinceAluminum may not be suitable for stable electrochemical sensing due tocorrosion, lithographic post-processing can be performed to replace orcoat it with a noble metal (e.g., platinum, gold, silver) for workingelectrodes and counter electrodes and Pt/PtOx (more stable referencethan Pt, easier to fabricate than Ag/AgCl) for reference electrodes.These post-processing steps are more specifically discussed below. Notethat the reference electrode can also be coated with a noble metalinstead of Pt/PtOx.

To directly detect glucose, working electrode(s) can be functionalizedwith GOx hydrogel (the chemistry and deposition of GOx hydrogel isdiscussed more below). In an embodiment, indirect glucose sensingthrough differential Oxygen sensing can be implemented by using oneworking electrode to measure background O₂ using a non-enzyme loadedhydrogel, while another working electrode is functionalized by GOxhydrogel to measure left-over Oxygen from the Glucose-Oxygen reaction(Oxygen consumed by the enzyme). The difference between these two Oxygenconcentrations can indicate the glucose concentration. The GOxfunctionalized electrodes can be intentionally placed apart (by havingWE, in between) to minimize crosstalk. On the electronics side, multiplepotentiostats (n) can be included to control the sensors. For example,GOx sensor working electrode can be held at +0.3V-0.6V with respect tothe reference electrode while an O₂ sensing working electrode can beheld at −0.3V-−0.5V with respect to the reference electrode (oxygendetection potential). The current from the potentiostats and thetemperature sensor can be digitized by the shared on-chip ADC in a timemultiplexed manner.

The strength of the sensing element signal can be proportional tosurface area of the electrode and the effective signal strength can beincreased by utilizing patterned or non-planar electrodes instead ofconventional planar electrodes. The working electrodes (and optionallycounter and reference electrodes) can be designed to have pillarstructures by patterning the top metal and passivation layers to enhancesensitivity as well as improving adhesion between the solid-state sensorand the hydrogel. Such structures can be formed using a semiconductorfabrication process, by post-processing or by a combination of both.Pillars created during a CMOS manufacturing process can have a size andspacing (determined by the CMOS process) of about 0.25 μm-25 μm andpreferably 2 μm-5 μm and height of about 0.1 μm-10 μm and preferably 2μm-5 μm (theoretically resulting in about 2-3× increase in sensorsignal). Post-processed pillar structures are discussed below. Bothfoundry and post-processed pillars share similar dimensions and can berounded or square. In accordance with some embodiments of thedisclosure, the pillars can be partially or completely etched to form amore planar structure. When completely etched, the pillars are totallyremoved. However, the resulting insulation structure on the surfaceresults in surface texturing after deposition of metal. When partiallyetched, pillars are thinned down and are shorter in height and possiblywidth.

An example of surface patterning which can be accomplished via foundrysemiconductor processes and a degree of post processing can be seen inFIG. 16A, FIG. 16B, and FIG. 16C. FIG. 16A, FIG. 16B, and FIG. 16C showhow pillar or hole structures can be made using semiconductor processeswith top insulation 122 generating spacing in top metal 121. The topinsulator can then be etched in post processing to create open areasbetween the pillar-like structures or the holes in the top surface.

In a hole-based pillar design, the electrodes are designed as metal meshand the hole is filled by the insulator. This method uses the smallfeatures size available in the CMOS process to form structures thatenable the formation of such high surface area structures with simplepostprocessing steps like wet etching.

In a pillar design, the metals are designed as separated electrodes withinsulator filled in the gap areas.

FIG. 16A shows the etching of top metal to create spacing. Afterwards,the pillar-like or hole containing structure can be coated with suitablemetal 125 (e.g., Platinum) using thin-film coating techniques (e.g.,electron beam evaporation, thermal evaporation, sputtering, or atomiclayer deposition).

FIG. 16B and FIG. 16C show two of the most common patterns that can beused to create the high surface area structures. FIG. 16B is beforeetching. FIG. 16C is after etching. After etching of either the topinsulator or top metal, these structures reveal reverse patterns of 3Dstructures (e.g., pillars and valleys or holes and top surface) that canthen be coated with suitable metal (e.g., Titanium and Platinum). InFIG. 16C, the structure on the left is called the pillar structure andthe structure on the right is called holes structure. It would beunderstood that in the holes structure, the bottom surface of the holesas a whole is typically less than a top surface of the structure.

The reverse-pillar or hole structure described above (and shown in FIG.18 ) is more robust than the pillar structure since it is made by holesinstead of free-standing pillars (shown in FIG. 17 , prone to defects asshown). In particular, FIG. 17 shows that the pillar structures can havedefects which manifest either as breakage of a set of pillars orcoverage of a set of pillars with some debris from other processes. Italso shows the position of interlayer vias 1291 on the bottom surface(valley area) of the electrode. FIG. 18 shows that the reverse-pillar orhole structure has much lower defects than the pillar structure. It alsoshows the position of interlayer vias 1291 on the bottom surface (bottomof each hole, bottom of some holes) of the electrode. Therefore, thereverse-pillar or holes structure provides better electrode integritywhile still providing high surface area. It also provides good hydrogeladhesion as the hydrogel fills the holes and forms a 3D structure withinthe electrode.

These patterned electrodes are different than those reported elsewhere.These electrodes are formed by using the design rules and materialsavailable in the CMOS process itself (e.g., insulator), instead ofhaving to form all the patterning on a silicon substrate afterwards.This simplifies the design and enables scalable and more controlledstructures. This also enables use of smaller features in advanced CMOSprocesses (e.g., 3 nm process) as those are achieved by advancedphotolithography methods not available in cleanrooms outside of theadvanced CMOS foundries. Moreover, this simplifies the post processingsteps which are otherwise difficult to match with the features andcontrol available in the CMOS fabrication process.

The size and shape of the electrode structure can be selected based uponthe sensing application and the desired integrated sensor circuit 1geometry. In accordance with some embodiments of the disclosure, thesensing element can include an arrangement of electrodes, e.g., acentrally located reference electrode (e.g., a rectangle of 50 μm by1500 μm), an outer counter electrode (e.g., a rectangle of 600 um by1500 μm), and a working electrode (e.g., a 150 μm by 1500 μm) locatedbetween the reference electrode and the counter electrode.

In another embodiment, the sensing element can include 2 workingelectrodes of 20 μm by 80 μm, 7 counter electrodes of 60 μm by 80 μm,and one reference electrode of 20 μm by 780 μm. In general, it would beacceptable to vary the area of the working electrode from half to doublethat listed directly previously. In various instances, it would beacceptable to make the working electrode with as little as 15× timessmaller area. The counter electrode should be always larger than theworking electrode. In various embodiments, the counter can be as littleas 3× the area of the working electrode; however, the maximum size ofthe counter electrode is only limited by the area of the implantablemonolithic integrated circuit. With respect to the reference electrode,smaller is better. For all practical purposes there is no electricallower bound on the size of the reference electrode. It needs to be closeto each counter. If, however, the reference electrode, is made of amaterial that may be consumed, such as AgCl, it is advisable for thecounter to be of similar size to the working electrode.

FIG. 10 shows embodiments of different electrode configurations of theworking 12, counter 13, and reference electrode 11 that can bemanufactured by various fabrication processes according to the presentdisclosure. This structure can be formed, for example, in the top metallayer by removing the top passivation layer of the chip to expose themetal sensor electrodes.

Monolithic Integrated Sensor Manufacturing

The process for forming the functionalized electrodes of the monolithicintegrated sensor circuit after receipt of a wafer from a commercialfoundry (e.g., TSMC) are now described in greater detail.

The wafer (diced or un-diced) as being processed is hereafter referredto as a monolithic circuit precursor. An example of a wafer can be seenat FIG. 14 step 1. FIG. 14 at step 1 is a side cross-section schematicview of an integrated sensor circuit precursor, showing a substrate 110,a metal-insulator-metal stack 129, a top metal 128, and a top insulatorwith an optional additional insulator/protection polymer layer 118.

Note that the top metal can be thick in high frequency CMOS processes.In some cases, a more suitable material can be coated on the top metalwithout etching it. For some other cases, a first step ofpost-processing involves removal of this top metal layer for bettercontrol of the morphology of the more suitable material. This etchingcan be achieved by using wet etching (e.g., using a mixture of Nitricacid and Phosphoric acid) or dry etching (e.g., Chlorine based RIEPlasma).

After receipt of a precursor of an implantable monolithic sensor circuitfrom a commercial foundry or after post-processing for removing, forexample, a thick top layer of top metal, lithographic (e.g.,photolithography) patterning can be done to expose the sensing elementelectrode while covering the rest of the wafer with a suitable material(e.g., photoresist). Note that, for some applications, this patterningcan be achieved using custom stencils, i.e., without lithography.

In FIG. 14 step 2 a precursor is patterned using lithography to create apatterned photoresist 148. Top first layer metal 128 of aMetal-Insulator-Metal (MIM) structure is Aluminum. Underneath the topmetal is the further Metal-Insulator-Metal (MIM) structure 129 found inCMOS devices. A silicon substrate is at 110. Note that severalelectrodes can be isolated from each other by top insulation 118 (Topinsulation in CMOS process is often a stack of Silicon Nitride layer ontop of Silicon oxide layer). Top insulation can be further augmented byan additional layer (e.g., Polyimide layer).

In photolithography, after mask deposition, the mask is etched. In FIG.14 step 3 a precursor is etched to remove the top metal 128 leaving thearea which was protected by photoresist 148 and further exposing asecond further layer of the Metal-Insulator-Metal (MIM) structure 129.The silicon substrate remains at 110. Electrodes remain isolated fromeach other by top insulation 118.

This patterning can be followed by deposition of a desired material(e.g., suitable metal stack); for example, a Ti (or Titanium-Tungsten)intermediate layer of small (e.g., 20 nm) thickness as the adhesionlayer (and to avoid corrosion of an underlying Aluminum layer) can bedeposited followed by deposition of relatively thicker (e.g., 100 nm) ofPlatinum (or any other noble metal or corrosion resistant conductivealloy). Physical vapor deposition (e.g., sputtering, e-beam deposition,and thermal evaporation), chemical vapor deposition, electroplating, andelectroless plating are different methods that can be used for thin filmdeposition. Sputtering can form a relatively rough surface compared toe-beam or thermal deposition both of which result in smootherelectrodes.

To achieve higher surface area and to enhance bonding between thesensing element and the subsequent chemistry layers, the metal surfacecan be designed to have rougher finish (as compared to smooth or mirrorfinish). This is achieved by controlling deposition method (e.g.,electron beam deposition, thermal evaporation, chemical vapordeposition, sputtering), depositional environment (e.g., pressure), anddeposition energy. In an embodiment, sputtering at 30 mTorr pressure and100 W DC power generates a metal coating with a high surface area for aplanar geometry.

FIG. 19 show a comparison between smooth Platinum coated using electronbeam deposition and low-pressure sputter deposition vs. high-pressuresputter deposition. It shows the smooth platinum layers of range 30-150nm (typical 100 nm) vs. rougher platinum layer of range 30-150 nm(typical 100 nm). A typical smooth layer of 100 nm thickness has rmsroughness in <3 nm range for 100 nm thickness while the rough platinumhas rms roughness >3 nm. Moreover, the rough Platinum appears to beporous under an SEM while the smooth Platinum appears more like acontinuous film.

A precursor is coated with a desired material (e.g., suitable metal) inFIG. 14 step 4. A thin layer of desired material (e.g., Platinum) 138 iscoated using a thin-film coating method. A substrate 110, ametal-insulator-metal stack 129 and a top insulator with an optionaladditional insulator/protection polymer layer 118 remains. The topinsulation electrode isolation region remains protected by photoresist148.

The use of patterned electrodes with one or more pillars can be helpfulas noted above because strength of the sensing element signal can beproportional to surface area of the electrode. As noted above patternedelectrodes can be formed by post-processing in an embodiment. Inpost-processing, it is possible to lithographically form the pattern forelectrodes in a lithographic material (e.g., photoresist) while coveringthe rest of the chip with the lithographic material. This is to be donebefore the deposition of a suitable metal stack. As an example, AZ 5214Eresist can be spun at 3000 rpm, baked at 95 degrees Celsius for 5minutes, and exposed using i-Line (e.g., 365 nm UV radiation) exposurein an MA6 mask aligner for 2 seconds. LOR resist can be used to helpwith liftoff. Image reversal can also be used for this purpose. In thiscase, a post-exposure bake at 110 degrees Celsius for 2 minutes isperformed followed by a flood exposure in MA6 for 3 seconds. For bothpositive and negative patterns, the resist can be developed in adeveloper (e.g., AZ300). This can be followed by sputtering of Ti (e.g.,20 nm) and/or TiW (20 nm) followed by Pt (100 nm). Sputtering parametersare optimized to achieve the desired morphology of the coated material(e.g., Pt). After sputtering, a conformal coating is achieved. SolventLift-off can then be performed (e.g., dipping sensors in acetone for 30minutes) to remove metal from unwanted areas and only keep those onsensing element electrodes. Alternatively, materials can first bedeposited everywhere and then etched with appropriate wet and/or dryetching methods.

A next step of the post-processing can be lift-off to remove metallayers from the unwanted regions. This is achieved by soaking the coateddevices in solvents. Alternatively, unwanted metals from coated devicescan be etched in appropriate solutions (e.g., in aqua regia).

In FIG. 14 step 5 a precursor is cleaned to remove the photoresist 148and excess deposited material and leave deposited material only indesired places (e.g., on integrated sensing element electrodes). Asubstrate 110, a metal-insulator-metal stack 129, a thin layer ofdesired material 138 and a top insulator with an optional additionalinsulator/protection polymer layer 118 remains.

It is possible to perform another lithography followed by silverdeposition, liftoff and Chlorine exposure through wet solution (e.g.,Ferric Chloride) or dry plasma (e.g., Chlorine Plasma) to createsilver-based reference electrodes (e.g., Ag/AgCl). Ag/AgCl referenceelectrodes are more suitable for some applications (e.g., open circuitpotential measurements).

It is also possible to create polymer structures around the sensingelement electrode area to create isolation or to improve chemicalfunctionalization. An example is the use of insulating material to coverthe unexposed parts of the top metal. This layer may consist of SiliconOxide and Silicon Nitride insulating layers found in standard CMOSprocess as well as additional insulating/polymer layers (e.g.,polyimide) to protect the underlying circuitry. Polymer walls around thesensor can be used to act as ‘well structures’ as well as ‘adhesionpromoting structure’ as some functionalization materials (e.g., SerumAlbumin based Hydrogel) adhere better to an activated polymer surfacethan to a Silicon Nitride insulation structure. In some cases, suchstructures can be provided by the CMOS foundry or a similar foundry aspart of the fabrication process. For example, polyimide structures canbe provided to the end-user by the CMOS foundry and can work as adhesionpromoters for some applications.

In accordance with some embodiments of the disclosure, the sensorelectrode surfaces can be activated (e.g., with glutaraldehyde or airplasma, oxygen plasma, or argon plasma) prior to functionalization layer(e.g., hydrogel) deposition. This activation can help with adhesion ofthe sensor chemistry with the sensor or the previously depositedchemistry layers. Surface structures and/or modifications can also actas grafts for a functionalization layer (e.g., hydrogel) and result in astronger adhesion and/or chemical interaction between the gel and thesensor electrodes.

Additionally, in optional embodiments of the disclosure, a layer thatcan limit sensor response to substances that interfere with sensoroperation can be applied to the surface of one or more of the electrodesbefore coating a functionalization layer (e.g., a hydrogel). Forexample, a layer of thin polymers (e.g., polyaniline) can be formed onthe sensor by spinning and UV/electron beam crosslinking. For example, alayer of poly-phenylenediamine polymer can be coated on electrodessurface using electrochemical deposition or UV crosslinking, before orafter the enzyme coating. This allows the sensor to not react toascorbic acid or acetaminophen which otherwise can create a false signalon platinum electrodes.

FIG. 14 step 6 demonstrates an embodiment of the use of finedroplet/spray deposition systems to precisely cover the integratedsensing element of a precursor with controlled amounts of afunctionalization layer. Optionally, the use of a fine droplet/spraydeposition system using solenoid and/or piezoelectric controlledactuation-based spray heads 188 simultaneously with spinning on table189 can improve the uniformity of the functionalization layer (e.g.,hydrogel). For example, solenoid and/or piezoelectric controlledactuation-based spray heads 188 can be used to deposit, in nanoliters,precise amounts of Glucose Oxidase solutions and crosslinking agents(e.g., Glutaraldehyde) solution droplets to make a glucose oxidase-basedhydrogel on the sensing electrode. The substrate can be spun after thedroplet deposition to control the thickness of the hydrogel layer moreprecisely. A substrate 110, a metal-insulator-metal stack 129, a thinlayer of desired material 138, and a top insulator with an optionaladditional insulator/protection polymer layer 118 remains.

An enzyme hydrogel is an exemplary functionalization layer. As notedabove, in one embodiment, an enzyme is immobilized on the sensingelement in a hydrogel (e.g., a cross-linked protein hydrogel). This canbe done at a thickness 0.01 μm to 50 μm. The enzyme hydrogel layer maybe under 3500 nm in thickness. Alternatively, the enzyme hydrogel layermay be less than 1000 nm in thickness. Further alternatively, the enzymehydrogel layer may be between 200 nm and 800 nm in thickness. Furtheralternatively, the enzyme hydrogel layer may be between 600 nm and 800nm in thickness.

This can be done using different techniques. As an example, this can bedone through immobilization of the enzyme such as GOx (Glucose Oxidase)in a hydrogel created by proteinaceous material with glutaraldehyde asthe crosslinking agent. The proteinaceous material can be a blockingagent such Human Serum Albumin (HSA) or Bovine Serum Albumin (BSA) orsome other Serum Albumin (SA). Herein a “blocking agent” is a materialthat blocks unwanted binding interactions of the sensor or sensorcomponents with tissue materials and fluids and avoids or decreasesfouling of the sensing element. Glutaraldehyde can be dispensed beforeapplication of the remaining elements of the hydrogel. Subsequently, amixture of GOx, Serum Albumin, and in some embodiments, catalase, can beplaced on the precursor. Glutaraldehyde can be used to aid hydrogelformation, and/or catalase can be used to increase sensor longevity bymitigating excess hydrogen peroxide production during glucose sensing.In accordance with some embodiments of the invention, it may bedesirable to remove excess hydrogen peroxide from the hydrogel duringglucose sensing, so a mixture of Catalase with GOx and Serum Albumin canbe used. In accordance with some embodiments of the disclosure, it maybe desirable to form the hydrogel after the solution is alreadydispensed on the electrode, by adding Glutaraldehyde to the mixtureafter it is dispensed on the electrode, for example, in a separate step.In accordance with some embodiments of the disclosure, GlucoseDehydrogenase can be used as the glucose sensing enzyme, in addition toor instead of Glucose Oxidase.

In accordance with some embodiments of the disclosure, to selectivelyfunctionalize the sensor electrodes, a precise deposition of nano- topico-liter of the hydrogel can be utilized. In one embodiment, thesubstrate can be heated or cooled and kept at a controlled temperature(e.g., 25 degrees Celsius to 35 degrees Celsius, with 25 degrees Celsiusbeing an embodiment) in a controlled environmental chamber (e.g., tocontrol temperature, humidity, chemical composition of the environment).Then, an accurate dispensing instrument (such as a BioJet Elite on aAD6020 aspirate dispense system by Biodot, Irvine, CA) with precise x,y, and z position control can be utilized. In one embodiment, thesprayed solutions are a protein solution of GOx and/or Catalase and HSA(1200 mg, 12 mg, and 1000 mg respectively in 15 ml DPBS, Sigma AldrichProduct codes G2133, SRE0041, SRP6182, D8537) and a crosslinking agentsolution of 1 w/w glutaraldehyde in DPBS (Sigma Aldrich, St. Louis, MO,product codes G5882, and D8537).

In accordance with some embodiments, deposition can be performed inthree steps to achieve a hydrogel of repeatable and controlled hardnessand composition: 1) dispensing glutaraldehyde, 2) dispensing the mixtureof GOx and SA, 3) dispensing glutaraldehyde. The three deposition stepscan be done almost simultaneously through the use of three dispensingnozzles as gel formation starts happening almost instantaneously once SAand glutaraldehyde come to contact. In a different method,glutaraldehyde is only dispensed once. With the three-step process, orwith a process where only steps 1 and 2 are performed, controlledtemperature (e.g., 25 degrees Celsius) of the sensing element electrodesurface and controlled environment (e.g., 80% RH, low particle count inair) during and after dispensing helps with uniform gel formation.

In accordance with a different embodiment of the disclosure, spincoating and/or spray coating can be used to achieve functionalization byapplying the sensing chemistry on the sensing elements, instead ofprecise deposition. In this method, enzyme hydrogel mixture is dispensedor sprayed on the precursor or even entire wafer using nano-dropletdispenser, spray head, or pipette. The hydrogel formulation can be thesame as that used in precise deposition. The wafer is then spun toachieve a thin sensing layer at controlled speed (between 200 to 20000rpm with 2000 rpm being an embodiment) for set time (10 seconds to 3minutes, another embodiment being 1 minute) to achieve a thin (10-50000nanometer thick, e.g., 2-6 micrometer thickness) layer sensingchemistry.

FIG. 15 shows one scheme of using spin-coating and/or spray coating topattern a stack of 3 (as an example, chemistries 191, 192, 193) coatingsuniformly across three electrodes on the sensing platform utilizing asacrificial layer (note the use of a sacrificial layer is not shown inFIG. 14 ). The chemistries can be deposited in fewer or more stepsdepending on what steps each chemistry can tolerate, and on thetreatments needed for each chemistry (e.g., plasma treatment, silanetreatment).

FIG. 15 Step 1 shows a post processed sensor chip, i.e., wafer withconductive material deposited upon sensing electrodes as seen in Step 5of FIG. 14 .

In FIG. 15 Step 2, a sacrificial layer (e.g., Polyacrylic acid) isdeposited on the sensor (e.g., by spin coating, spray coating). In someembodiments a water-soluble sacrificial layer can be used. For example,a mixture of Polyacrylic acid can be used to make a sacrificial layer.Same can be done by using Polyvinyl alcohol or other water-solublematerials as well as material that can be dissolved in other solvents.In some embodiments (not pictured) water-soluble sacrificial layers canbe used to protect the sensing chemistries to allow for patterning thesensing chemistries with a photoresist via a lift-off process, where thewater-soluble layer would protect the enzyme layer from the photoresistand chemical and physical methods used to pattern and clean thephotoresist, including but not limited to developer, acetone, or plasmaetching.

In FIG. 15 Step 3, a patterning chemistry 181 (e.g., photoresist, e-beamresist) is deposited over the sacrificial layer 182 (e.g., by spincoating, spray coating).

In FIG. 15 Step 4 the patterning chemistry 181 (e.g., photoresist,e-beam resist) is patterned.

In FIG. 15 Step 5, after patterning of the patterning chemistry 181 inStep 4, the sacrificial layer 182 is dissolved in a manner as tocorrespond to the pattern created by the patterning chemistry.

In FIG. 15 Step 6, the patterning chemistry 181 is completely removed.

In FIG. 15 Step 7, chemistry 191 is deposited by spin coating.

In FIG. 15 Step 8, chemistry 192 is deposited by spin coating.

In FIG. 15 Step 9, chemistry 193 is deposited by spin coating.

In FIG. 15 Step 10, sacrificial layer 182 is dissolved leaving thetarget chemistries 191, 192, and 193 only on the target electrodes.

In accordance with a further alternative embodiment, instead of precisedeposition or spin coating and/or spray coating, the functionalizationlayer (e.g., hydrogel: crosslinking agent and or the protein mixtures)can be deposited on the wafer via dipping. In some instances, the sensorchips or the entire wafer can be mounted on a substrate that can bedipped vertically or horizontally in a solution of enzyme or enzymes andserum albumin and optionally glutaraldehyde. The hydrogel formulationcan be the same as that used in precise deposition. The substrate can bedipped and dried one or more times for a total processing time rangingfrom 2 minutes to 2 hours depending on desired gel thickness andconsistency. In some embodiments, the sensors can be dipped for oneminute and dried in a chamber with 80% relative humidity for 5 minutesfor 10 cycles for a total processing time of 60 minutes. In accordancewith some embodiments, the sensors can be dipped in protein solutionsand glutaraldehyde solution, sequentially. For instance, if there are avariety of sensing chemistries dispensed on the sensor, and many ofthese produce hydrogen peroxide, then, after the coating(s) aredispensed, the whole wafer can be dip coated in catalase solutionfollowed by dip coating in glutaraldehyde to immobilize the catalase onthe sensors' surface. In some embodiments of the disclosure, a cleaningsolution such as DPBS can be used between dipping steps in order toprevent beading of the solutions on the sensors and resulting loss ofuniformity.

In some cases, the hydrogel can conform to the pattern of underlyingpillars to result in patterned layers. Such patterning can allow forfaster response time (lesser delay) but may have a shorter lifetimecompared to sensors with thicker hydrogels and polymer layers. Thehydrogel and polymer can be shaped either like blocks of materialscovering the pillars or shaped liked pillars with empty space in betweenthe pillars. The different shapes can be obtained by process control(e.g., if droplet functionalization of hydrogel is used a block ofhydrogel formed; however, if spray, a thin layer mostly sticks topillars and provides a conformal coating).

In accordance with some embodiments of the disclosure, each workingelectrode can be isolated from the rest of the working electrodes andallow for unique functionalization of individual electrodes, e.g.,different electrode sensing chemistries. These methods for selectivefunctionalization of individual working electrodes are in addition tothe ability of precise deposition to achieve this result as noted above.Options to facilitate selective functionalization include stencils,lithographic patterning, nanoimprint lithography, and selectiveactivation. In accordance with some embodiments of the disclosure, whereisolation is required, all the sensing element components for any oneanalyte application can be dedicated (e.g., separate working, referenceand counter electrodes) and isolated from others.

Stencils can be used to selectively functionalize sensing elements withdifferent chemistries. In these embodiments, a stencil, e.g., a metalsheet with holes corresponding to sensing element surfaces, can beplaced on the die or wafer. Then sensing chemistries can be dispensed,dropped, dipped, or sprayed, or otherwise deposited. In some embodimentsspraying is used. Then the stencil can be lifted from the surface toleave defined sensing chemistries deposited on sensors. The stencilprocess can be repeated or combined with other processes to achieve avariety of chemistries.

Alternatively, wafer scale lithographic patterning can be used. In someof these embodiments, a light-active chemical (e.g., a photoresist) canbe placed on the die or wafer and patterned using light and a developeras known to those skilled in the art. Then dispensing, spin or spraycoating, dipping, or any method described in the above surfacefunctionalization paragraphs herein can be employed to deposit sensorchemistries on the specific sensors.

Nanoimprint lithography is yet another technique that can be used forthis purpose. In this case, a special printing head/stamp can be used totransfer small gels on to the sensing element surfaces (e.g., electrodesurfaces). The gel is first formed on this stamp (which can be madeusing lithographic patterning or molding) using any of the methodsdiscussed herein (e.g., nano-droplet dispensing, spin coating, spraycoating, dipping). Then the stamp is placed on the desired wafer and amethod is used to release the hydrogel to the specific sensors on thewafer. This is facilitated either by increasing gel adhesion with thesensors on the wafer (e.g., by surface activation of sensors andparticularly surfaces of sensing elements in a manner such as withoxygen argon or air plasma) or by using heat/UV to create some change onthe stamp which releases the gel.

Specific sensing elements can also be patterned by selectivelyactivating the sensing element surfaces (e.g., with an oxygen, argon, orair plasma, or chemical modification) and sensing chemistries can bedeposited using any of the methods discussed herein (e.g., nano-dropletdispensing, spin coating, spray coating, dipping). Then, the sensingchemistries can be removed (e.g., washed with deionized water, or amixture of deionized water and detergent such as 10% (w/w) Extran(MilliporeSigma, Burlington, MA) in deionized water) such that onlysensing chemistries bonded to the activated surfaces remain.

FIG. 14 step 7 shows an embodiment of a device after deposition of thefunctionalization layer 158 using solenoid and/or piezoelectriccontrolled actuation-based spray heads. A substrate 110, ametal-insulator-metal stack 129, a thin layer of desired material 138,and a top insulator with an optional additional insulator/protectionpolymer layer 118 remains.

FIG. 14 step 8 shows an embodiment of the use of spray coating (e.g.,using fine dispensing heads 186) to make a film (stack of one or morethin films) on sensor surface to coat the functionalization layer, usingan appropriate spray head 186 which showers microdroplets on an area ofthe sensing element. A substrate 110, a metal-insulator-metal stack 129,a thin layer of desired material 138, a functionalization layer 158 anda top insulator with an optional additional insulator/protection polymerlayer 118 remains. FIG. 27 similarly shows how a soft material likehydrogel 190 is coated on the sensors using a thin-film coating process(e.g., spin coating, spraying).

In an optional embodiment, a drying step can be done in a chambersaturated with crosslinking agent vapor, e.g., glutaraldehyde vapor, toaid or obviate the need for crosslinking via crosslinking agent viafurther application. For example, for vapor crosslinking a crosslinkingagent in the solution may not be required. In accordance with someembodiments of the disclosure, the protein solutions can be preciselydeposited (using a precision instrument as described above) on thesensor electrodes and spread using spinning in the presence ofcrosslinking agent vapor.

Before and/or after functionalization, (but typically afterfunctionalization as noted in FIG. 14 ) different films (e.g., membranematerials) can be used to protect and/or restrain the functionalizationmaterials on the sensing element 160 and achieve a desirable signalresponse for a particular sensor configuration. In some embodiments ofthe disclosure, a diffusion limiting layer can be useful.

For example, in the body there is 30 to 300 times more Glucose thanOxygen. If the sensing mechanism has a 1:1 stoichiometry (e.g., Glucosedetection using GOx uses 1 molecule of Oxygen for every molecule ofGlucose), then the sensor placed without a limiting membrane will belimited by oxygen concentration and will not be able to sense glucosefor the entire physiological concentration (e.g., 40-400 mg/dl). Apolymer membrane can be deposited to act as a suitable diffusion barrierthat allows oxygen to go through unhindered but hinders glucosediffusion.

Examples of workable polymer membrane materials include polyurethane, amixture of polyurethane and silicone, as well as a mixture ofpolyurethane and PEG. In accordance with some embodiments, the thicknessof the polymer membrane can be in the range from 0.1 micron to 15microns. The polymer membrane may be between 200 nm and 10,500 nm thick.Alternatively, the polymer membrane may be between 200 nm and 1500 nmthick.

Adhesion between the membrane coating and the underlying hydrogel, orbetween layers of coating, can be facilitated by use of chemicals (e.g.,silanes, aldehydes) and/or physical processes (e.g., corona treatment,oxygen plasma, gas plasma, mechanical roughening). Specific membranematerials and construction can be used to further improve sensorperformance. In one embodiment of the disclosure, a combinationcomposition of polyurethane and silicone can act as a filter to regulatediffusion of glucose and as an oxygen recycling membrane as well asproviding a biocompatible material. Oxygen recycling can improve theefficacy of the sensor, while the biocompatibility can allow the sensorto work for longer. To cover the sensor uniformly and minimize sensor tosensor and batch to batch variation, membranes can be deposited on thesensor through spotting (droplet coating), spraying or throughwafer-level spin coating. Membranes can also be deposited on thebackside of the wafer to increase biocompatibility. Another method touniformly deposit membranes is to employ spray coating with a specialinstrument utilizing overlap between multiple depositions to achieve auniform overall thickness.

A specific workable polyurethane membrane coating process includesloading 1% PurSil from DSM in THF (DSM Biomedical, Exton, PA and SigmaAldrich, St. Louis MO) into an Air-jet spray coating unit (BioDot,Irvine, CA). A single coat of 1.25 microliter/cm is applied at 9 PSIpressure on sensor area, with dispensing height and aperture optimizedfor each coating unit installation. The wafer is dried in a vacuum ovenat 35 degrees Celsius and 25.6 mm-Hg pressure for an hour and in ambientconditions for at least 12 hours (overnight). A second coat is applied,and sensors are dried with the same parameters. The sensors are allowedto stabilize in PBS (Sigma Aldrich, St. Louis MO) for 72 hours andcharacterized for analyte response.

Optionally, a membrane coating can also be patterned to reduce cellattachment. This patterning can be done using oxygen plasma or usingnanoimprint lithography (bio-stamping). For oxygen plasma, after themembrane is coated on the surface, it is exposed to a high-power oxygenplasma (e.g., 300 Watt) in a plasma chamber without any mask or with alithographic mask (e.g., AZ5214E for Photolithography, PMMA for electronbeam lithography) or a stencil. The oxygen plasma will etch the exposedmaterial and the material under the etch mask will be protected, henceshaping the surface. For stamping, the membrane is coated and patternedwhile being on a different substrate (e.g., Silicon substrate). Next,the membrane and the sensor surfaces are exposed to some process thatprepares their surface for strong adhesion. One example is to exposeboth materials to a short, low-power oxygen plasma dose or to chemicallinkers like Xylene. Next, the membrane is placed on the sensor surfaceand the bonding process is allowed to continue for some time. At theend, the substrate is gently removed, and the patterned membrane remainson the sensor surface.

FIG. 14 step 9 shows an embodiment of a device after deposition of asecond film 168, such as a interference rejection layers, immuneresponse suppressing layer, and/or biocompatibility layer, on thesensing element. A substrate 110, a metal-insulator-metal stack 129, athin layer of desired material 138, a functionalization layer 158 and atop insulator with an optional additional insulator/protection polymerlayer 118 remains.

Another example of polymer coating, other than a membrane, is use ofinterference rejection layers that can be coated on the electrodesbefore or after surface functionalization. These layers can similarly becoated using spraying, dip coating, electrochemical coating, spin and/orspray coating. In accordance with some embodiments of the invention, acoating including o-phenylenediamine can be used for rejecting Ascorbicacid and/or Acetaminophen in glucose sensing applications.

Another example of polymer coating which can be used includes immuneresponse suppressing layers. Implanted sensors can be attacked by theforeign body or immune response of the body. This can be mitigated byincorporating coatings that inhibit response and/or mitigate the effectsand decrease this foreign body response. Drugs such as dexamethasone ornitric oxide limit such response. Note that, in some embodiments of thedisclosure, drugs that inhibit adverse response by the body (e.g.,dexamethasone, nitric oxide) can be mixed, encapsulated, or chemicallyincluded in the functionalization layers and/or membrane layers, insteadof a separate coating, in a way that allows slow release of the drugsthroughout the functional lifetime of the sensor. Another example ofpolymer coating which can be used includes a biocompatibility layer.

To improve biocompatibility of the system, the sensor can be coated witha biocompatible material. Proteins attach to hydrophobic surfaces; thus,an option to improve biocompatibility is to cover the device with ahydrophilic or superhydrophilic polymer (Note that polyurethane, anoption for membrane formation, is hydrophobic). For example, thebiocompatible material can be poly-HEMA. A layer of pHEMA can be formedwith a thickness of 5 microns to 100 microns and a preferable thicknessof about microns. In some cases, a copolymer of a biocompatible materialcan be made with polyurethane to coat the device in a single step.

An additional/alternative biocompatibility layer can be the depositionof Titanium and/or Platinum or a catalase hydrogel to mitigate theeffects of reactive oxygen species. For example, a layer of catalase canbe coated over top a layer of GOx hydrogel. The layer of catalase canbe, for example 0.05 μm to 25 μm, alternatively 0.1 μm to 3 μm. It hasbeen shown that reactive oxygen species from glucose oxidase can damagesurrounding tissue. Platinum is known to breakdown the reactive speciesinto less corrosive byproducts and hence is ideal for this application.For instance, platinum microspheres can be dispersed within the pHEMAhydrogel or any other hydrogel and used to coat the surface. In short, asolution of platinum spheres in water can be used to make the hydrogeland ultrasonic mixing of the components can be used to ensure properdispersion of the spheres throughout the hydrogel. In a furtherembodiment of the disclosure both a hydrophilic biocompatibility layerand a means to quench reactive oxygen species are used in combinationwith an immune response limiting element. For example, dexamethasone(0.01%-3% w/w) can be mixed into the polyurethane layer. PolyHEMA layercan be patterned with nanoimprinting to achieve a super-hydrophilicsurface. 2 nm thick Ti/2 nm thick Pt can then be sputtered on thesurface to quench reactive oxygen species while maintaining superhydrophilicity and porosity. As noted above, multiple layers of membraneand/or polymer materials can be applied to the implantable monolithicsensing circuit.

Another example of biocompatible layer (also known as sensor-tissueinterface layer) is PVA. A specific workable PVA includes the following:Poly(vinyl alcohol) (MW 89,00-98,000 99%+ Hydrolyzed, Sigma 341584) wasmixed with DI water 4% (w/v) (1-10% acceptable), by slowly adding PVAinto DI water heated to 80 C and stirred at 400 rpm with a stir bar. Thesolution was capped and let stir for 12 hours prior to use.

To make the PVA layer, the solution is drop casted or spin coated on thesensor surface. In a typical recipe, the solution is dispensed to covermore than 90% of the surface to be coated (e.g., wafer, die, or flex PCBwith the sensors attached). The surface is then spun, first at a lowerspeed (e.g., 500 rpm) for 10 seconds to spread the solution and thenaccelerated to a higher speed (e.g., 500-5000 rpm with 3000 rpm as atypical case) to get to a desired thickness. The sensors are then leftto sit in a controlled environment (e.g., an oven set at 35° C.) for aknown time (e.g., 24 hours) for the PVA gel to form.

FIG. 14 step 10 shows an instance of how dipping can be used to apply asmall volute of material (e.g., polymer 178) on the sidewalls. Asubstrate 110, a metal-insulator-metal stack 129, a thin layer ofdesired material 138, a functionalization layer 158, a film 168 and atop insulator with an optional additional insulator/protection polymerlayer 118 remains. FIG. 27 shows how another material can be coated onthe pre-coated sensors using a thin-film coating process (e.g., dipcoating, spraying).

After deposition of one or more membrane and/or polymer materials, in anembodiment, the implantable monolithic sensing circuit can be consideredcomplete. FIG. 2B shows the components of an embodiment of a completelyprocessed implantable monolithic integrated sensor circuit including thesilicon substrate with integrated contact pad 175 attached to flexibleconnector, the integrated working electrodes, WE1, WE2, and WE3,representing the integrated sensing element, said working electrodesseparated by insulation walls 118, with the working electrodes coatedwith the functional matrix example Gox-loaded hydrogel 158 and thepolymer coating example PU membrane 168.

In accordance with some embodiments of the disclosure, a post-processedsensor wafer can be cleaned with deionized water and/or pressurized gasand dried in vacuum oven (20-400 degrees Celsius, e.g., 40-200 Celsius;0 to 30 mm-Hg below atmosphere, e.g., 26 mm-Hg). In accordance with someembodiments of the disclosure, a cleaning and drying step can befollowed by a plasma cleaning and surface activation step. In someembodiments, the sensor can be cleaned under 50-600 mTorr pressure ofoxygen or air or argon plasma with a power of 75-400 W. In someembodiments, Oxygen plasma at 100-500 mTorr, with a power of 90-200 Wcan be used. In accordance with some embodiments of the disclosure,after post-processing and drying, wafers or sensors can be placed in ahumidity controlled nanoliter dispenser equipped with an aluminumchilled plate calibrated to be able to operate at 80% RH and 25 degreesCelsius plate temperature.

Flexible Connector

In the present disclosure, a physical intermediate transdermal component3 is used to connect the external transmitter 2 located on the skin tothe monolithic integrated sensor circuit 1 in the analyte concentrationmeasurement system.

One typical embodiment of the intermediate component is a bidirectionalconnector. The connector can be a single wire, two wire, three wire,four wire, or higher instance flexible connector 3. In some embodiments,the connector can be a flexible printed circuit board with the wiresbeing conductive trace(s) therein.

Wires/conductive trace(s) can be made of copper, aluminum, gold, orother conductive material. The connector can comprise a biocompatiblepolymer such as parylene-C, liquid crystal polymers (LCP), or polyimide,which may almost completely cover any wire/conductive trace(s) exceptfor points necessary to connect to the external transmitter 2 orintegrated sensor circuit 1.

FIG. 20 shows several schemes to connect the integrated sensor circuit 1with the transmitter 2 via a flexible connector 3 comprising varying thenumbers of wires/conductive traces. The connector 3 can have 1 ormultiple wires/conductive traces to flow power and data between thesensor 1 and the transmitter 2.

Data communication can be performed using a two-wire connector per“communication over power” through superimposition of the data signalover the power wire. The same conductors can be used for both data andpower (e.g., to save space) as shown in the 2-wire configurations shownin the top of the FIG. 20 . One of the wires can act as the power supplyand signal and the other as the ground. If data and power share aconducting line, a coupling-decoupling (Co-De) network (also termed asMultiplexing-Demultiplexing or MUX/DEMUX network) on both sides of thesame conductor is to be used. The Co-De network can be designed toseparate data from power on either or both sides.

With regard to MUX/DEMUX network, at the transmitting side, this can bedone through an AC coupling capacitor and a tri-state driver in oneparticular embodiment. At the receiving end, an AC decoupling capacitorcan be utilized to extract data by decoupling it from the power signalfollowed by a hysteresis comparator which detects the signal while beingresilient to the signal noise. Since both transmitter and the monolithicintegrated circuit will be operating in both transmitting and receivingmodes, the aforementioned circuitries can be incorporated in both thetransmitter and the sensor. Since the data is modulated over the powersignal, the power signal received at the sensor is not clean. Hence avoltage regulator is utilized at the sensor to create a clean and stableDC supply. This DC power can be sent to the potentiostat which powers upthe integrated sensing element. If a read command is received from thetransmitter, the monolithic sensor circuit can send digital readings viathe two-wire flexible connector to the transmitter.

However, different conductors for data and power can also be used tosimplify the system design as shown in the 3-wire configuration of FIG.20 . If full-duplex communication is desired, separate conductors can beextended to use n number of conductors, e.g., for sending redox voltagefrom the transmitter to the sensor, for sending data for multiplesensors from the sensor to the transmitter and so on.

The flexible connector can fit in a 16 gauge to 32 gauge needle with 23to 28 gauge being an alternative range and 26 gauge being a furtheralternative. Rectangular needles can be used instead of standardhypodermic needles, in which case an equivalent needle gauge descriptioncan be used as the standard needle gauges are defined for cylindricaltubes and the needles made from those tubes. The equivalent needle gaugecan be defined by comparing the hypotenuse of the rectangular needle(based upon the triangle formed by the base and the wall height) withthe diameter of the cylindrical needle.

With regard to the mechanical properties of the connector, saidproperties of the connector 3 can be controlled by controlling itsconductive materials, insulating layers, and shape. It is important tomanage the mechanical properties of the connector to provide a connectorof sufficient rigidity and strength to be successfully applied by anapplicator yet flexible enough as to not induce a strong immune responseby damaging surrounding tissue during user movement.

For example, stiffeners (e.g., polyimide, glass, steel, FR4) can beattached as a bottom layer of the flexible connector 3 to provideadditional rigidity.

Alternatively, a layer of bottom Cu can be used to increase thestiffness of the panel, as shown in FIG. 26 . This use of standardflexible PCB insulated copper traces to control the stiffness of theflexible connector is a unique feature of this design. Control on sizeand thickness (e.g., 0.25 Oz, 0.5 Oz, 1 Oz, 2 Oz etc. Cu) can provide acontrol on the mechanical properties of the panel. For example, a panelwith a dummy bottom metal underneath the top metal designs can bestronger than a panel without the bottom Cu. The Cu is completelycovered with sealed insulators to prevent any interaction with the bodyfluids. In some cases, Cu can be covered with more biocompatiblematerials (e.g., noble metals like Palladium and Gold) or can even bereplaced by those. This stiffness control enables thinner devices ascompared to the use of stiffeners, and hence enables the use of smallerneedle sizes which reduces insertion pain and foreign body response. Thetop metal and the bottom metal layers may also be referred to as “firstmetal layer” and “second metal layer”, respectively.

With regard to a connector (e.g., flex-PCB), the total thickness in someembodiments can be 100 to 250 μm, preferably 150 μm. A flex-PCB can bemade of three layers, five layers, or more depending on the design ofthe conductive traces and stiffness requirements.

In a specific exemplary embodiment, the flexible connector 3 is aflexible PCB. The flexible PCB is of five total layers (generally knownas 2-layer PCB based upon 2 conductive layers) with a top Cu layer of ½oz (18 μm) and bottom Cu layer of 2 oz (70 μm). A top overlay can beused of 1.5 mils (37 μm)−1 mil adhesive+0.5 mils. The substrate of thePCB can be for example 1 mil (25 μm). Preferably, a flex-PCB willinvolve a stiffener in areas where mechanical strength is requires,e.g., its connection with other electronics (e.g., the transmitterboard). The stiffener can be for example 75 μm (3 mils of FR4) or asimilar thickness of stainless steel or Polyimide.

The bonding pads of the flexible connector 3 can be made with differentfinishes including Electroless Nickel, Immersion Gold (ENIG),Electroless Nickel, Electroless Palladium, Immersion Gold (ENEPIG),Immersion Gold on Cu directly as few examples. The pads are used on oneend to connect with the integrated sensor circuit 1 and on the other endwith the transmitter 2 using different methods detailed next.

A variety of processes can be used to join the connector to themonolithic sensor circuit. In accordance with some embodiments of thedisclosure, a two-wire connector 3 is made up of a flexible polymersubstrate, e.g., polyimide or parylene-C, with thin copper traces (e.g.,20-100 μm wide, alternatively 20-50 μm wide; and 20-100 μm apart,alternatively preferably 20-50 μm apart; with 1-30 μm thickness,alternatively 10-30 μm thickness). The copper traces are sandwiched bythe polymer substrate to avoid exposure to bodily fluids. On theimplantable monolithic integrated circuit 1 side two bonding pads arecreated by removing a passivation layer during a CMOS manufacturingprocess, removing any biocompatible membrane and/or polymer and platinggold on top of the copper wire connection on the circuit. At theconnection site to the implantable monolithic integrated circuit 1, thecopper traces can be exposed and covered with gold for good connectionto the sensor using a flip-chip bonding technique or wire bonding.

FIG. 25 provides a packaging process scheme of the sensor circuit 1 ontothe connector 3. FIG. 25 Step 1 illustrates a sensor circuit 1positioned adjacent to a connector 3 (e.g., a flexible PCB panel). Theconnector 3 may include conductive traces 302 in a middle layer of theflexible PCB panel, i.e., under a PCB top layer 303. The sensor circuit1 can include multiple sensors 160 at some distance from CMOS pads 175.FIG. 25 Step 2 shows the bonding of the sensor circuit 1 to the flexibleconnector at 301 through die attach, followed by forming conductivetraces between the sensor circuit 301 and flexible connector 3 via CMOSpads 175 and wire bonds 314. The wire bonds can be made using manymetals like Aluminum, Gold, Copper, Platinum etc. The methods to makewire bonds include thermosonic bonding which uses a combination of heatand ultrasonic power to attach thin (e.g., 25 μm diameter) wire on CMOSpads and flexible connector pads. In one example, 25 μm gold wire isused to make connections between the CMOS pads and the flexibleconnector pads. An image of a wire bonded sensor circuit 1 can be seenbelow.

FIG. 26 Step 3 shows the conductive pads and the conducting interface onthe sensor circuit 1 and the flexible connector array 301 protected forwater ingress and mechanical abrasion by covering those with abiocompatible insulating material 313. An example is a biocompatibleinsulating material (e.g., 31CL) that is deposited in small volumes(e.g., 1-20 microliters) on the wire bonds followed by thermal curing(e.g., at 100° C. for 1 hour). Sometimes, the encapsulation process isdone in multiple steps to ensure complete coverage of the wire bonds. Animage of a wire bonded sensor circuit 1 can be seen below.

FIG. 27 shows an instance of how an entire row of implantable monolithicintegrated circuits 610 can be attached to a flexible PCB 620 using wirebonding 630. In either flip-chip or wire bonding, the connection betweenthe connector 3 and the implantable monolithic can be hermeticallysealed with a biocompatible polymer (e.g., parylene-C). Smartnano-dispensing robots (e.g., Nordson EFD) can be used to preciselydispense appropriate biocompatible polymers (e.g., for hermetic sealing)like 31CL (Henkel Loctite EA M-31CL Medical Device Epoxy) in small andcontrolled volumes (e.g., 1-20 microliters, typical 2 microliters). FIG.27 shows an example of how an insulating material 640 can be used tohermetically seal the sensor-PCB interface. One or more soft materialslike enzyme hydrogel 190 can be coated on the sensors using a thin-filmcoating process (e.g., spin coating, spraying). This can be followed bycoating a second soft material like polyurethane 189. The integratedsensor 1 and flexible PCB can then be diced together.

Another technique to integrate the sensor 1 with the flexible connector2 is the use of flip-chip bonding. FIG. 28 provides an example offlip-chip bonding. In this method, the sensor 1 pads 175 and theflexible connector pads 302 are aligned and pressed together. A softmaterial (e.g, Au, InSn) bump 1751 is optionally used to fuse the twopads together via a combination of heat, pressure, and optionallyultrasonic energy. Conductive bump 1751 sits on CMOS pad 175. Afterfusion, the contact pads are covered with insulting epoxy (e.g., usingunderfill). As seen in FIG. 29 , the interface between the flex PCB andthe CMOS pad 175 can be covered in insulating epoxy 313.

Another method can use flip-chip bonding using anisotropic conductiveadhesive (ACA). This can be accomplished through bonding the monolithicintegrated sensor circuit 1 to the flexible connector using ACA. Ifdesired, conductive bumps are formed on the CMOS pads, followed by ACAcoating on the CMOS pads and/or the Flexible connector pads. Afterwards,the CMOS and Flex connector pads are aligned and pressed together for acertain duration at controlled temperature and pressure. This forms anelectrical connection in one direction (e.g., vertical) between desiredpads on CMOS and the flexible PCB. The interface between the flex PCBand the CMOS pad can be covered in insulating material, e.g., using anunderfill or a similar process. Additional pads can be used on the CMOSand corresponding pads on the flexible PCB for a stronger and moresymmetric bonding. An optional cutout in the flexible PCB can ensurethat the sensor area 160 is exposed to the environment for sensing.

Another technique to integrate the sensor 1 with the flexible connector3 is to embed the sensor within the flexible connector, as shown in FIG.30 . In this case, the sensor chip 1 is placed between the flexibleconnector layers. Electrical connection between the sensor chip 1 andthe flexible connector 3 is made by using interlay conductors (likevias) 38 in the flexible connector. Flexible PCB contact pads are at 39.The active sensor area 160 is kept open by patterning the dielectriclayers of the flexible connector as shown in FIG. 30 . Functionalmaterial 19 can be deposited over active sensor area 160. This methodenables a completely solid-state device with minimal height and can bescaled rapidly.

FIG. 31 shows another scheme of connecting the monolithic integratedsensor circuit 1 and the flexible connector 3 (e.g., flexible PCB). Inthis scheme, the sensor 1 uses through-silicon-via (TSV) technology tohave contact pads 19 under it. In this scheme, a trench is etched fromthe backside of the CMOS sensor circuit 1 silicon substrate 110 to theconductive trace using a combination of wet and/or dry etching methods.Then the sidewall of the silicon is covered with and insulator andfinally with a conductive material that runs from the backside of theCMOS sensor circuit to the front side, e.g., to chip pad 175. These padson the backside of the CMOS sensor circuit 1 can be attached to theflexible PCB conductive trace 302 using different methods (e.g., bumpbonding, conducting adhesive like Anisotropic conductive adhesive orACA) depending upon the size and the application. Next, the conductinginterface is covered with an insulating material 313.

Yet another embodiment for packaging together the monolithic integratedcircuit 1 and the flexible PCB is provided in FIG. 32 . In this case,the monolithic integrated sensor circuit 1 is placed inside a well inthe flexible connector 3. The gap between the sensor circuit pads 175and the flexible connector pads 302 is filled with an insulatingmaterial 313, followed by forming the connection between the two pads314 (e.g., using wire bonding, or conducting epoxy, etc.) followed bycovering it with a thin layer of insulating material 313.

Moving to connecting the connector 3 to the transmitter 2, variousoptions exist. The 2-wire flexible connector can be soldered to thetransmitter circuit board. Alternatively, it can be soldered to amedical grade connector which can then be connected to a connector onthe transmitter. In a separate embodiment, micro connectors can be usedto connect the connector 3 to the transmitter 2. In addition toconnecting the wires of the connector to the implantable monolithicsensor circuit, in various embodiments the implantable monolithic sensorcircuit 1 can be further secured to the connector 3.

FIG. 33 shows a scheme of using the integrated sensor circuit 1 on oneside of the flexible connector and using the other side of the flexibleconnector as a secondary sensor 9 (e.g., a 2 electrode or a 3 electrodesensor design). The flexible connector 3 on the other far side uses aconnector that can connect with its conducting traces on both sides ofthe flexible connector, thus enabling connection with both theintegrated sensor circuit 1 and the secondary sensor 9. The secondarysensor can be used for validation and control purposes. It is read bythe transmitter circuit which in a simple embodiment would just read itsvoltage (e.g., to confirm if the sensor is inserted under the skin, toconfirm if the sensor is being perfused). In a different embodiment, thetransmitter would use a potentiostat circuit to read the electrochemicalsignal from secondary sensor 9. The secondary sensor 9 can befunctionalized as is the integrated sensor circuit 1 above. Thissecondary sensor 9 can then be compared with the reading from theintegrated sensor 1 for reliability check of the platform as well as forother similar checks (e.g., to see if sensor is being properlyperfused).

FIG. 34 shows another embodiment of the system architecture in which theflexible connector 3 not only has conductors for the electricalconnections but also has conductors working as a sensor/sensorelectrodes, e.g., electrochemical sensor 40. The flexible connector 3has contact pads 42 that are connected with the sensor 40 via traces 41.Moreover, an application-specific integrated circuit (ASIC) 30 is placedin close proximity to the sensor to help process its data. The ASICcomprises at least a potentiostat. The contact pads 42 can be connectedwith the contact pads 31 on the ASIC 30. The ASIC 30 contains anotherset of contact pads 32 for connection with the conductors 302 of theflexible connector 3. The connection interface is covered with aninsulator.

FIG. 34 Step 1 shows the ASIC attached to the flexible assembly. FIG. 34Step 2 shows the connection between the contact pads 42 on the flexibleconnector and the pads 31 of the ASIC and between conductor 302 and thecontact pad 32 of the ASIC via bonding wires 43 and 44 respectively.FIG. 34 Step 3 shows both exposed connection interfaces can be coveredwith insulators 45 and 46.

This scheme enables the passive CGM designs to benefit from mostfeatures of the integrated sensor 1 design (e.g., close proximity of thesensor and the electronics) without having to change the way existingmanufacturers, like Medtronic, manufacture their electrochemical sensors40.

Transmitter

The analyte concentration measurement system requires an externaltransmitter 2 located on top of the skin of the user to secure thetransdermal connector 3 and the monolithic sensing circuit 1. Thetransmitter 2 can provide electrical power to the monolithic integratedcircuit 1 via the flexible connector 3. The external transmitter 2 canbe used to power monolithic integrated circuit 1 and to communicate withimplantable monolithic integrated circuit 1, i.e., to receive data fromand send data (e.g., commands) to sensor before and after the monolithicintegrated circuit 1 is inserted.

The transmitter is an electronic device with a power source (e.g., abattery). A printed circuit board (PCB) can be used to make atransmitter system with all components to function as the transmitter 2.One or both of thick (rigid) and thin (flexible) PCB technologies can beused, depending upon the application.

In various embodiments, the transmitter 2 can also communicate with themonolithic integrated circuit 1 via the flexible connector 3. Thecommunication can be bi-directional or unidirectional, whereinoptionally bi-directional communication is sequential, meaning thatfirst the transmitter sends a command (e.g., a tag) to the monolithicintegrated circuit 1 to trigger analyte measurement by the sensor. Afterthe transmitter sends this tag, it can go into receiving mode (i.e., itwaits for the sensor to send it the measured data). Error correctionschemes can be employed to minimize the error in this communication.Parity-bit based designs and more advanced error correction codes can beused as well. Different types of modulation schemes can be employed forthis communication.

A schematic diagram of an example of a transmitter according to someembodiments of the disclosure can be seen in FIG. 35 . The transmitter 2can include a transceiver 210 that can generate and detect communicationsignals flowing over a combined power/data connection through wire 3, ahigh-performance digital microprocessor unit 211 for system control, alow-energy wireless communication chipset 213 (e.g., Bluetooth), anantenna 218 for wireless communication (e.g., Bluetooth antenna), apower management unit 214 connected to a battery 212, and a voltageregulator unit 217 to provide regulated power to the implantablemonolithic integrated sensor circuit 1.

In an exemplary commercial off the shelf transmitter design, asystem-on-chip can be used to provide the functions of a microprocessor,a wireless communication link (e.g., BLE) with an integrated antenna, aswell as on-system memory. Similarly, a battery unit can contain thebattery along with a battery management circuit. Examples of a suitablesystem-on-chip for use in the transmitter include TAIYO YUDEN® EYSHSNZWZBLUETOOTH® Low Energy Module and Lilypad coin cell battery module.Additional electronic circuit like a transceiver can be implemented topre-process the incoming data from the sensor circuit 1 to make iteasier to be read by the processor. In one example, the transceiver mayinclude operational amplifiers (e.g., Texas Instrument TLV9061) toamplify the incoming signal and a comparator (e.g., Texas InstrumentTLV7011) to create a rail-to-rail signal. In some examples, thistransceiver circuit can be implemented within the SoC e.g., by usingon-chip amplifiers and comparators.

Although a transmitter could be implemented via a commercial off theshelf technology, it is possible to use a custom solution usingapplication specific integrated circuits (ASIC). FIG. 36 shows adetailed electrical schematic of a transmitter embodiment which can befully integrated into a custom ASIC. It shows a battery 212 powers thesystem via a voltage regulator 217 that uses a voltage reference 216 toset the output voltage. Different types of batteries can be used topower the system, including Lithium-ion, Silver, Zinc, Lithium polymer,thin-film etc. For example, a lithium-ion battery voltage can range from3.3V to 2.8V based upon remaining charge. The voltage regulator canprovide a stable output voltage (e.g., 3V) independent of such batterystatus, thus helping with stable circuit operation for longer duration.The voltage regulatory powers the rest of the transmitter componentsincluding a microprocessor/microcontroller 211, a wireless communicationtransceiver (e.g., BLE transceiver) 213 feeding an antenna 218. Themicrocontroller generates a suitable tag signal to communicate with thereceiver (implantable monolithic sensing circuit 1). The tag signal isfed to a TX unit 215 that shapes the tag and couples it to the wiredconnection 3 via a DC decoupling capacitor 242. A switch 228 is closedwhen this tag signal is being transmitter while another switch 229 iskept open to prevent leakage of this signal into the Rx unit 210 of thetransmitter. After the tag is sent and a certain duration is passed, asignal is expected from the sensor 2. At this time, the switch 228 isopened while the switch 229 feeding the signal to the Rx unit 210 oftransmitter 2 is closed. The Rx unit 210 can shape the received signal,e.g., by using a monostable circuit to increase the width of receivedsignal. The received signal passes through the DC decoupling capacitor242 before being fed to the Rx unit 210. The main voltage regulator 217feeds power to a secondary voltage regulator 231 that is used togenerate a more suitable voltage (e.g., 2V) for the receiver (e.g.,implantable monolithic sensing circuit 1) using voltage referencecircuit 232. The resistor 243 is used to isolate the voltage regulatorfrom the transmission line 3. The resistor 243 and the capacitor 242 areexample embodiments and can be replaced with other types of electricalnetworks to perform this same function of combining the data with thepower on one end (from transmitter to the sensor) and separating thedata and power on the other end (from sensor to the transmitter).

The housing options for the PCB of the transmitter varies. FIG. 37 showsan implementation of the transmitter. It includes the connector 3attached (e.g., by wire bonding, soldering, through a medical grademicro-connector, etc.) to the transmitter which further consists of aprinted circuit board 21 housing the battery 212, regulator 217, andintegrated BLE transceiver and application controller/processorSystem-on-chip 250. This part represents integration of severalcomponents (e.g., microprocessor, transceiver, antenna). The transmitteris housed in a sealed plastic/rubber housing 22 for environmentalprotection and easy handling during usage. FIG. 37B shows images furtherinclusive of connectors 32 and 34, which can mate to connect the printedcircuit board and connector 3.

FIG. 38 shows another embodiment of the implementation of thetransmitter 2 using a printed circuit board with electronic componentsand a power source, housed within a suitable casing. The connector 3 onone hand attaches to the transmitter circuit and on the other hand tothe integrated sensor circuit 1. In embodiments, the connection betweenthe transmitter circuit and connector 3 is off to a side of thetransmitter assembly to enable more compact assembly.

The transmitter can wirelessly communicate the data to a hub or smartdevice (e.g., a phone, a tablet, or a special separate device 4). Inaccordance with some embodiments, the hub or smart device can beconnected (either by wire or wirelessly) to a cloud server via a network(e.g., the Internet, a private network such as virtual private network(VPN), or a public network). The transmitter uses a low-power wirelesscommunication technology (e.g., Bluetooth, Zigbee) to communicate withthe smart device. The transmitter may use a standard BLE profile (e.g.,the CGM profile) to enable a simple interface with other devices (e.g.,to form artificial pancreas). A secure BLE connection can be stablishedbetween the transmitter and the host device using techniques likededicated digital keys.

The transmitter is programmed using a firmware to enable operationsdesired in a typical implementation. The firmware is designed to performthe operations necessary to control the operation of the transmitterhardware to match with system needs. The firmware is stored on apermanent memory block on the transmitter. The memory typically resideswithin the Bluetooth module of the transmitter. Many BLE modules havethe option to wake-up the system using an NFC link. The transmitterfirmware is designed to keep it in low power mode until any such wake-upcalls are received. This can enable the transmitter to remain in lowpower mode till the user is ready to operate the unit (e.g., aftersensor insertion). Then the transmitter can be taken out of the sleepmode to start a secure dedicated connection with a desired reader andstart communicating the data.

After such wake-up, the transmitter firmware (designed as a BLE client)looks for a BLE master to connect to. If such a connection (device) isavailable, the transmitter establishes a connection to the device (e.g.,smart reader). After this, the transmitter tests the available powerlevel and upon validation of a stable power supply, it looks to see if asuitable sensor is attached (via the amount of current draw). It alsochecks sensor readings to be within normal range with a normal rate ofchange (as programmed in its memory during manufacturing) to confirm ifthe sensor is operating properly (e.g., in warm-up period). After thewarm-up period is complete (prep-programmed during manufacturing), thetransmitter checks the reading values and the rates of change todetermine if the sensor is operating properly.

Since the sensor is a smart CMOS device, the data coming from it isdifferent than the data coming from the passive CGMs in otherapplications (just the sensor current). In this case, the header(preamble), the footer contains the sensor batch and type and arecompared against the IDs stored on the transmitter to confirm that thesensor and the transmitter are indeed matching. Moreover, the errordetection schemes (e.g., CRC) can quickly tell if the received data isproperly encoded. The data from the sensors, including the temperaturesensor can then be compared against each other and the factory storedvalues (e.g., in a look-up table) to confirm if the sensor is completelyoperational and is properly inserted in the tissue. After this, thetransmitter sends a tag to the sensor telling it to start sending thesensor data at a desired interval (e.g., once every 2 seconds). Thisdata is then read by the transmitter, tested, and combined (e.g.,averaged over 3 sensors) before sending it to the reader at a certainfrequency (e.g., once/minute).

The transmitter can also include usability features like a temperaturesensor to detect environment temperature reading which can then becompared with the temperature sensor on the integrated sensor 1 todetect change in temperature from the environment to the tissue aroundthe sensor. Another optional embodiment is for the transmitter to have alight sensor (e.g., a Photodiode) that acts as a way for the system todetect its state (i.e., still in the packaging, vs outside of thepackaging).

System Packaging and Operation

The sensing platform can be loaded inside an applicator device and theentire assembly can be sterilized (e.g., by Synergy Health (San Diego,CA)). To sterilize the device before embedding it inside the body,conventional methods of sterilization (e.g., steam, Ethylene Oxide) canbe utilized. In one particular embodiment, Electron-beam (e-beam)sterilization can be used to sterilize the sensor as well as theapplicator once the sensor is pre-loaded in the applicator. Theunderlying electronics can be designed to be resilient to e-beamradiation. The enzyme chemistry can be characterized to calibrate forany changes in the enzyme chemistry response due to sterilization. Inone embodiment, 25 kGray of e-beam irradiation can be sufficient tosterilize the sensor without impeding its function. Sensors can beplaced inside the applicator and then the whole assembly can besterilized. In a different embodiment, the integrated sensor 1 and theconnector 3 assembly, the transmitter 2, and the applicator can besterilized separately using different methods (e.g., sensor 1 andconnector 3 assembly using e-beam, transmitter 2 and applicator usingethylene oxide) followed by sterile assembly.

One or more sensors 1 can be placed in desired tissue locations using anapplicator as noted above. As noted throughout the present disclosure,in embodiments, the implantable monolithic integrated sensor circuit 1can be used to measure glucose levels in the user. The readout procedurefor collecting glucose data from the implantable monolithic integratedcircuit 1 can start with energizing the implantable monolithicintegrated circuit 1 through transmission of power signal (through wiredconnection) from the external transmitter 2. The external transmitter 2can be configured to select the appropriate powering mode(continuous/intermittent) based on user or clinician input.

The external transmitter 2 can receive sensor data, display sensor data,store the data, relay it to a smart device 4, or send it a remoteserver. External transmitter 2, smart device/communication device, orremote server can relay and process the sensor data in a mannercommensurate with its processing, storage, or battery capability. Thedata processed in external transmitter 2, smart device/communicationdevice, or remote server can be relayed to external transmitter 2, smartdevice/communication device, or remote server to provide, display, orstore, information (e.g., blood glucose levels, daily trends) orpredictions thereof or suggestions (e.g., behavioral changes, treatmentchanges) based on sensor data or predictions.

After the sensor assembly is removed from the packaging, the transmitter2 needs to be turned on to power the sensor 1 and to communicate withthe reader 4. Different schemes can be used for this purpose. Forexample, a photosensitive detector (e.g., a photodiode) can be used todetect opening of the package. This signal can then be used to turn onthe transmitter power supply and start the initialization and connectionprocedure.

In a different embodiment, a wake-up-in-the-field method using an NFCpairing capability embedded in the transmitter can be used to turn onthe transmitter operation after the user opens the sensor packaging. Thetransmitter is paired with an NFC enabled device to send it the turn-onsignal. The transmitter then starts the initialization process.

FIG. 39 shows an algorithmic scheme to program the processor (e.g., amicrocontroller) in the transmitter (which acts as the brain of thetransmitter) with a firmware to control its operation. It shows that thetransmitter (microcontroller) is programmed to stay in a low power(e.g., deep sleep) state. After the user opens the packaging andperforms a turn-on operation (e.g., either by using an NFC device topair with it) or the transmitter automatically detects the user's intentto use it (e.g., by detecting a change in background conditions likelight via a photodiode), it performs an initialization sequence whichincludes a self-test as well as a scan for the reader via BLE. Once itfinds a matching reader, it connects with it via BLE. Next, it tests ifa good sensor is connected by performing electrical measurements (e.g.,voltage drop, current draw) and by sending a command signal and testingthe response. After that, it starts transmitting the tag signal andstart reading the corresponding data from the sensor chip. Thetransmitter separates the power and data signals (via mux/demux) andsends it to the microcontroller which preprocesses the data (e.g., checkfor proper preamble, proper data coding scheme, packet length, packetduration), checks if it detects any error (via errordetection/correction code), and separates the data from all 4 (3electrochemical and 1 temperature) sensors. It then sends the data tothe reader via BLE.

Multi-Sensor Signal Processing Schemes

As described above in relation to FIGS. 4 and 5 , the integrated sensingelement 160 includes three (or more) working electrodes, each of whichgenerates a respective sensing data signal. The quantities and qualitiesof various parameters of interest can be determined as a function ofthese data signals. In some embodiments of the disclosure, the datasignals can be combined with other reference and/or stored data signalsto generate the quantity and/or quality of parameters of interest. In anexample, the three (or more) working electrodes are used to detect theconcentration of the same analyte (e.g., glucose).

The presented device has a unique advantage of having multiple sensingelectrodes (e.g., multiple working electrodes) on a single device. Thedata from these multiple electrodes can help improve the quality of thesensor data. Different approaches can be used to do so and generallyreferred to as data fusion techniques. For example, in one method, thesignal from all 3 working electrodes can be averaged to minimize theeffect of noise or skewness from the data. In another case, the medianof all readings can be used to do so. In a different method, thereadings from all 3 working electrodes can be compared to eliminate anyoutliers (e.g., if one electrode reading is significantly different thanthe other two). This method is called voting. In case of outliers,interpolation can be used to fill in the gap of the missing outlierdata. Or that sensor's data can be discarded in decision making for aparticular duration.

In general, the sensor reading at any given time is a function of thereading from all the working electrodes (referred to as “sensors” belowfor brevity). For example, for the case of 3 glucose sensors, the sensoroutput can be written as a function of the 3 sensors' readings asg=ƒ(g ¹ ,g ² ,g ³)The function can be of many types. In a simple way, an arithmeticfunction (e.g., weighted average) can be defined based upon the sensorvalues

$g = {\sum\limits_{k = 1}^{n}{a^{k}g^{k}}}$

In a simple case, the equation can becomeg=a ¹ g ¹ +a ² g ² +a ³ g ³

The above equation describes a weighted averaging technique to convertthe data from 3 sensors into a single data stream. The coefficients (a¹,a², a³) can be calculated during testing of the sensors in known glucoseconcentrations in the lab and in test subjects, i.e., in calibration(done at the factory, in the field, or in a combination). The sensorvalues (g¹, g², g³) are readings of the 3 sensors, respectively.

A control logic (e.g., a processor, a processing unit, a controller, ora microcontroller etc.) may be employed to implement the function f. Thecontrol logic may be part of the monolithic integrated sensor circuit 1.For example, the control logic may be part of the sensor signalacquisition unit 130. Alternatively, the control logic may beincorporated within the transmitter 2, and for example be part of theprocessor 211 (FIG. 35 ). Further alternatively, the control logic maybe external to the transmitter 2 and the monolithic integrated sensorcircuit 1. In an example, the control logic may be part of thesmartphone 4 (FIG. 1 ) or part of a computer which processes the data ofthe secure database.

In real-time use, several different approaches can be used to takeadvantage from the multiple sensors. For example, a mathematicalfunction (e.g., arithmetic averaging) can be applied to all sensors andany sensor with large deviation from the result (e.g., mean value) isconsidered to have erroneous reading. Also, an average of a subset ofsensors (e.g., 2 sensors out of 3) can be taken and compared against thereading from the sensors. If any sensor has large deviation from thisaverage, the sensor is considered to have erroneous reading. Thisprocess is repeated for all sensor combinations. At the end, any sensorwith an unacceptably large error (its combinations have largest error)is considered to have erroneous reading at that point, which means itsreading is eliminated (its weighting coefficient becomes 0). FIG. 40shows a scheme for using the multiple on-chip sensors (workingelectrodes) data as well as previous readings and a personalized (foreach patient) patient-sensor model to generate the best outcome (glucosevalue) at a given time. The scheme uses the current data from allsensors (e.g., 3 electrochemical and 1 temperature) and uses that tocalculate errors among the sensor values. It also compares the readingswith the previous reading to decide if the new values arephysiologically accurate. By comparing the sensor readings amongthemselves and the errors, the system decided if a sensor hasunacceptable level of error. In that case, it discards that sensors anduses rest of the sensors data to generate a weighted average as thecurrent value of the sensor.

TABLE 1 T(m) Ref S1 S2 S3 Wavg M Avg 1-2 Avg 1-3 Avg 2-3 E1 E 2 E3 M ESME DS BS AE EWD 0 10.05 10 10.1 8 9.37 10 10.05 9 9.05 0 0.1 2 2 3 310.05 0 0.68 5 10.3 10.2 10.5 8.2 9.63 10.2 10.35 9.2 9.35 0 0.3 2 2 3 310.35 0.05 0.67 10 10.3 10.1 10.3 9.1 9.83 10.1 10.2 9.6 9.7 0 0.2 1 1 33 10.2 0.1 0.47 15 9.6 9.5 9.75 10 9.75 9.75 9.625 9.75 9.875 0.25 00.25 0.25 3 3 9.625 0.025 0.15 20 10.2 10.1 10.4 15 11.83 10.4 10.2512.55 12.7 0.3 0 4.6 4.6 3 3 10.25 0.05 1.63 25 13.38 13.5 13.3 19 15.2713.5 13.4 16.25 16.15 0 0.2 5.5 5.5 3 3 13.4 0.02 1.89 30 15.45 15.215.4 32 20.87 15.4 15.3 23.6 23.7 0.2 0 16.6 16.6 3 3 15.3 0.15 5.42 3518 17.8 17.4 21 18.73 17.8 17.6 19.4 19.2 0 0.4 3.2 3.2 3 3 17.6 0.40.73 40 19.3 19.9 19.5 25 21.47 19.9 19.7 22.45 22.25 0 0.4 5.1 5.1 3 319.7 0.4 2.17 45 21.5 21.75 21.9 18 20.55 21.75 21.825 19.875 19.95 00.15 3.75 3.75 3 3 21.825 0.325 0.95 50 23.05 23.1 23.9 22.9 23.30 23.123.5 23 23.4 0 0.8 0.2 0.8 2 2 23 0.05 0.25 55 22.5 22.3 22.9 22.5 22.5722.5 22.6 22.4 22.7 0.2 0.4 0 0.4 2 2 22.4 0.1 0.07 60 20.25 20.3 20.121 20.47 20.3 20.2 20.65 20.55 0 0.2 0.7 0.7 3 3 20.2 0.05 0.22 65 19.5519.1 19.6 19.5 19.40 19.5 19.35 19.3 19.55 0.4 0.1 0 0.4 1 1 19.55 00.15 70 18.4 18.5 18.9 18.4 18.60 18.5 18.7 18.45 18.65 0 0.4 0.1 0.4 22 18.45 0.05 0.20 1.77 15.64

Table 1 above provides an example of fusing sensor data from threesensors presented by the present disclosure. Table 1 is based uponEXAMPLE 6 below. Each of the three sensors is used to detect theconcentration of glucose within a patient. The column “T(m)” indicatesthe relative time at which the sensor readings were taken. It can beseen that the sensor readings were taken once every 5 minutes. Thecolumn “Ref” indicates a glucose reading of the same patient using aDEXCOM® sensor. The columns “1”, “2” and “3” indicate the glucosereadings of the three sensors presented by the present disclosure,respectively. The column “Wavg” indicates the weighted average value ofthe glucose readings of the three sensors, with the weights being 1 foreach sensor in this example. In other words, weighted average “Wavg” isequal to (S1+S2+S3)/3. The column “M” indicates the median value of theglucose readings of the three sensors. The median value is the middlevalue of the glucose readings of the three sensors, which is higher thanthe lowest glucose reading and lower than the highest glucose reading.For example, at time ‘0’, the median value is the glucose reading ofsensor 1, and at time ‘55’, the median value is the glucose reading ofsensor 3. The column “Avg 1-2” indicates the average value of theglucose readings of Sensor 1 and Sensor 2. Similarly, the columns “Avg1-3” and “Avg 2-3” indicate the average value of the glucose readings ofSensor 1 and Sensor 3, and the average value of the glucose readings ofSensor 2 and Sensor 3, respectively. The column “E1” indicates theabstract value of the difference between the glucose reading of Sensor 1and the Median value. Similarly, the columns “E2” and “E3” indicates theabstract value of the difference between the glucose reading of Sensor 2and the Median value, and the abstract value of the difference betweenthe glucose reading of Sensor 3 and the Median value, respectively. Thecolumn “ME” indicates the maximum error, i.e., the highest value amongstE1, E2 and E3, and the column “SME” indicates which sensor has themaximum error value. For example, at time ‘0’, it is Sensor 3 which hasthe maximum error value with respect to the median value. The column“DS” indicates the identity of the sensor of which the glucose readingis to be discarded. In this example, the identity of the sensor of whichthe glucose reading is to be discarded is the same as the sensor whichhas the maximum error value. For example, at time ‘0’, the glucosereading of Sensor 3 is to be discarded. The column “BS” (‘best signal’)indicates the output glucose reading (i.e., the fused sensor data),which is generated as an average of the glucose readings of theremaining two sensors. For example, at time ‘0’, the value BS is equalto Avg 1-2, because the glucose reading of Sensor 3 is discarded.

The columns “AE” (‘actual error’) is used to assess the quality of thefused sensor data. In this example, the actual error indicates theabstract value of the difference between the best signal “BS” (i.e., thefused sensor data) and the reference “Ref”. In generating the bestsignal “BS”, a sensor signal which is significantly different than theother two sensor signals (e.g., an outlier) has been eliminated. Theelimination of the outlier has significantly improved the accuracy ofthe fused sensor data. This is evident by comparing the column “AE” andthe column “EWD” (‘error without discarding worst sensor’). The column“EWD” indicates the abstract value of the difference between theweighted average “Wavg” of the three sensor readings and the reference“Ref”. It can be seen that when the worst sensor reading (i.e., outlier)is not discarded, the fused sensor data produces an accumulated errorwhich is 8.8 times the accumulated error generated when the worst sensorreading is discarded.

Since the sensor can provide a reading much faster (e.g., 10milliseconds) than the reporting rate (e.g., 1 minute), differentfiltering schemes (e.g., averaging) can be used in time domain also toremove some of the noise. Different types of filters (e.g., Kalmanfilters) can be used to filter the data to remove noise. The filtercoefficients can initially be assigned based upon sensor design andtesting data from. The coefficients can be updated after CGM data isavailable (e.g., from a training model or from the same patient viareference sources) to improve its functionality.

Since not only the recent data but the rate of change of glucose (owingto several last readings) is available, an estimate of the next readingcan also be calculated based upon these readings using differentestimation techniques like least square estimation (LSE). This estimatecan be compared against the reading from the sensors and the differenceis used to see if the sensor is reading an expected or an erroneousreading. Based upon the difference, a confidence level can be assignedto the reading of each sensor. Furthermore, based upon the rate ofchange, a prediction level can be assigned to hypo/hyper glycemiaoccurrence, and the user can be alerted.

The data from the on-chip sensors can be used to test system state,integrity, and operation. For example, the typical readings from theelectrochemical sensors can be used to check if the sensor was properlyinserted in the body. Since the current range of the sensor in differentconditions, i.e., in the air, in saline, in the skin, etc. are known,the value of the sensor current after possible insertion can be used tocheck whether the sensor is being properly wetted from the tissue fluid.

In a data analysis method, the next reading may be predicted based upona number of previous readings and a patient-sensor model. The patientsensor model may indicate, for instance, the daily variation of theglucose concentration of a particular patient. The confidence level ofthe best reading may be determined based upon a comparison between thepredicted next reading and the fused reading (i.e., best signaldescribed above).

Note that a control logic may carry out processing steps as shown inFIG. 40 and/or other data analysis processes above.

Temperature Effects

The glucose sensor readings can be further improved by calibrating thosewith the local temperature measured by the on-chip temperature sensor(shown in FIGS. 4 and 5 ). In general, the instantaneous glucosereadings can be represented by the following relationship between thesensor current (i(t)), the calibration coefficients (a,b), referencebody temperature (e.g., 37° C.), and the instantaneous local temperature(T(t)),g(t)=f(i(t),a,b,T(t))

For example, one example of the relationship can be (for temperature indegree Celsius)g(t)=(a*i(t)+b)*(T _(B) /T(t)))

The exact relationship is determined by testing a subset of devices froma manufacturing batch and by fitting the relationship between the sensorcurrent and glucose reading. The calibration coefficients are adjustedto represent the decrease in sensor current in the body as compared tothe in vitro testing.

The following table shows that by calibrating the glucose sensor readingwith the local temperature data, the error of the final, calibrated,glucose reading can be reduced significantly.

Within the monolithic integrated sensor circuit 1, the temperaturesensor can be located within 10 microns of the sensing element 160, suchthat the output of the temperature sensor and the output of the sensingelement 160 are measured from roughly the same portion of human tissue.Alternatively, the temperature sensor can be within 5 or 2 millimetersof the entirety of all working electrodes. Alternatively, thetemperature sensor is connected within 0.5 millimeter or 400 microns tothe entirety of the working electrodes.

These arrangements improve the accuracy of the calibrated glucosereading, as compared to an arrangement where the temperature sensor andthe sensing element 160 are sufficiently away from one another.

Current Glucose Reduction in error in mg/dl by (nA) (mg/dl) temperaturesensor 30 130 20 50 285 85

The temperature sensor data can also be used for system, operation, andsignal integrity purposes. For example, since the body temperature has adefined range (e.g., 37° C.+−2° C.), the temperature sensor readingafter insertion can be used to determine if the insertion was successfulas shown in FIG. 58B. Together with electrochemical sensor signal changein the body, this can provide a multi-modal and hence more reliable thana single-sensor scheme to confirm successful sensor insertion in thetissue. Since there would be a temperature sensor available in thetransmitter, the gradient of temperature reading between the two canalso be a good indicator of the successful insertion as well as ofsensor depth. In extreme environments, this gradient can be used tocalibrate the sensor if it moves too far from the normal bodytemperature.

In some cases when other sources of temperature variations are minimal,hypoglycemia can also be detected by a drop in the peripheraltemperature. Hence, a drop in the on-chip temperature sensor without asignificant drop in the environment temperature (e.g., measured by anexternal temperature sensor like one in the transmitter) can be anindicative of hypoglycemia. Such effects are documented in our humanstudies as seen in the EXAMPLES and further show the value of theon-chip temperature sensor.

An increase in environment temperature can also cause an increase insubcutaneous tissue temperature which can increase the enzymaticactivity which in turn can increase sensor current. This can makedetecting hypoglycemia harder in a hot environment. This can bemitigated by using the temperature sensor which can calibrate the sensorresponse according to the surrounding temperature and hence can mitigatethis risk substantially as compared to other sensors without suchtemperature sensor. An opposite effect (considering higher glucose ashypoglycemia) can occur in colder environments. The integratedtemperature sensor can avoid this problem as well.

A control logic (e.g., a processor, a processing unit, a controller, ora microcontroller etc.) may be used to process (or calibrate) theinformation from the sensing element 160 by taking into account thetemperature measurement obtain from the temperature sensor. The controllogic may be part of the monolithic integrated sensor circuit 1. Forexample, the control logic may be the control logic of the powermanagement unit 120, or may be part of the sensor signal acquisitionunit 130. Alternatively, the control logic may be incorporated withinthe transmitter 2, and be part of the processor 211 (FIG. 35 ). Furtheralternatively, the control logic may be external to the transmitter 2and the monolithic integrated sensor circuit 1. In an example, thecontrol logic may be part of the smartphone 4 (FIG. 1 ) or part of acomputer which processes the data of the secure database 5 (FIG. 1 ).

Sensor Calibration

After sensor fabrication, the calibration schemes are tested and storedeither on the sensor memory or the transmitter memory or on both. Themanufacturing batch (e.g., a wafer or a batch of wafers) is tested inknown environments (e.g., known temperature range, known glucose range)by following a sampling scheme (e.g., 10% sensors). The results from thetesting scheme are mathematically analyzed to determine a relationshipbetween sensor response and the sensor outputs. This relationship iscalled the calibration algorithm and is stored so that the system canconvert the response from the other (remaining 90%) sensors intocalibrated values of temperature and glucose. The variations among thesensors in the test sample lead to some variations in the calibratedoutput. The results can be further improved by doing a personalizedcalibration by a user. After the system is used for each user, anypotential calibration results are stored in the user profile and can beused to personalize the calibration algorithm for that person in future.Moreover, this data along with any demographic data shared by the usercan be used to further optimize the calibration algorithm for that usesas well as for that particular demographic population.

For temperature sensor calibration, the sensors are tested in a salinetest solution for body temperature range (e.g., 35-41° C.). For glucosesensor calibration, the sensors are calibrated in physiologicalconcentration (e.g., 40-400 mg/dl). The data is fed to the mathematicalmodel formed during the clinical trials that formalizes the relationshipbetween the in-vitro performance and the in-vivo performance of thesensor. The model generates calibration coefficients which are thenstored in system memory (e.g., typically in the transmitter but can bedone on the monolithic integrated sensor 1 or on both).

Sensor Application

One or more sensors 1 can be placed in desired tissue locations using anapplicator. In a simple embodiment, the applicator consists of a plasticbody to hold the transmitter and a needle to hold the sensor as well asto pierce the skin. The transmitter and sensor assembly can be placedinside the injector from the top side by using a removable top cover ofthe injector body. Once the sensor is in place, the top cover can be putback. The whole assembly can then be placed on the skin (away from majororgans, preferably on the arm, thighs, or the belly) and pressed down.This inserts the needle under the skin and allows analyte transfer tothe sensor.

FIG. 41 shows how the monolithic integrated circuit, wire, andtransmitter are placed inside an injector system consisting of anexternal plastic body 81, adhesive patch 82, and needle 83. FIG. 34 alsoshows the placement of the sensor and injector assembly under the skinwhen the system is pushed down. This allows body analyte to reach thesensor surface and the system to start monitoring the concentration ofone or more analyte in the body.

However, such a simple applicator suffers from several disadvantages.Specifically, needle 83 is left in place which can irritate surroundingtissue. Accordingly, in a more preferred embodiment an applicator isused which leaves the integrated sensor 1 in place while itself beingcompletely removed from the user.

Such an applicator 8 is designed to place the integrated sensor 1connected with the connector 3 at desired depth and position under theskin (e.g., 1-10 mm under the skin, straight or at an angle) wherein theconnector 3 is attached to the transmitter 2. An exemplary embodimentcan be seen in FIG. 42 .

The applicator 8 uses a cylindrical design consisting of 11 components.The external body 81, internal compression spring 812-813, chassis 82, atriggering arm 83, needle holder 84, needle compression spring 85,hypodermic needle 86, transmitter (also known as antenna) holder 87, andtransmitter assembly 21. The push-retract mechanism of this applicatoris initiated through the external body 81 in which two cylinders, height24.6 mm and radius 3.38 mm diameter, that mate through the chassis 82and onto the antenna holder 87. The internal compression springs 812-813are 25.4 mm long, have a maximum load of 1.05 lbs, and are grade 302stainless steel. Once the push mechanism is complete, the retractionmechanism will begin by releasing the triggering arm 83 and allowingneedle compression spring 85 to decompress and retract the needle intothe compartment at the base of the external body 81. This will allow theuser to avoid any injuries by safely capturing the needle and allowingfor it to be a onetime inject device.

The applicator 8 comprises an external cover 81 (plastic body to enabletactile holding of the applicator on the skin). The external cover canbe shaped as a cylinder with the bottom posterior side forming an opencircular aperture and a top anterior side being closed, wherein the sidecurved surface may include a major groove 89 to allow fingers of a userto more easily grab the applicator 8.

Inside the applicator lies a triggering guide 82. It is a cylindricalshape with a slot 823 in it to receive the two bars of the triggeringarm 83 (explained below). The external cover 81 contains multipleprotrusions 88 extending towards the triggering guide 82 andlongitudinally from the interior of the anterior top towards the bottomposterior open circular aperture. The protrusions 88 assist alignment ofthe triggering guide 82 during assembly and prevent the assembly guide82 from contacting the interior top anterior surface of the externalcover 81 during operation. The external cover also includes two rods 810extending from the interior of the anterior top towards the bottomposterior open circular aperture. The rods 810 function as guide polesfor springs 812 wherein the rods are configured to have an exteriordiameter smaller than the mean diameter of the respective spring. Therods can extend from bases which can be similarly shaped to protrusions88 and can operate as bases for the tip of the coil of springs 812 tosit thereon.

A triggering arm 83 passes through triggering guide 82. The triggeringarm 83 can comprise a disc with an open central aperture extending fromthe top and bottom surface of the disc as well as two bars extendingfrom opposite sides of the disc. The triggering arm 83 can guide needleholder 84 into the appropriate position based upon the status of thetrigger mechanism as the two bars extending from opposite sides of thetriggering arm 83. The triggering arm 83 is further in contact with aspring 85 which is used to store a spring force for activation of thetrigger mechanism. The needle 86 is held inside the needle holder 84 viathe spring 85 and is attached to an assembly holder 87. The far end ofthe assembly holder is also used to hold the external transmitter 81inside the applicator. The transmitter 2 is kept inside the transmitterassembly holder 87. The applicator components are shown in FIG. 42 .

The system is designed to fit and move smoothly via groves and matchingpatterns on different parts. Briefly, the external cover 81 has threegroves 814 to fit the assembly holder 87 via three features 874. It alsohas two springs 812 and 813 to provide smooth movement of the assembly.The triggering guide 82 has corresponding holes 821 and 822 to matchwith the springs 812 and 813 of the cover 81. Additionally, it has threecut-outs 824 to allow sliding movement of the assembly holder 87 viathree holders 872 on the assembly holder 87. It also has 4 slots 825 toenable alignment of assembly holder 87 with it (triggering guide 82).The needle holder 84 has a holding structure 841 that enables it toslide with the assembly holder 87 within the groves 824 of thetriggering guide 82.

During the initial stage, the applicator is in the armed state (as shownin FIG. 44 ) by the assembly process. Briefly, the assembly processstarts with placing the assembly holder 87 into the triggering guide 82.Next, the spring 85 is put on the needle side of the needle holder 84and the triggering arm 83 is placed on top of needle holder 84. Next,the triggering arm is pushed down and rotated clockwise to lock thetriggering arm in the armed state. The entire assembly is then fittedinside the external cover. Now the spring is loaded and the needle 86 isbelow the transmitter assembly holder 87. The sensor 1 lies inside theneedle 86.

The assembly of the applicator involved mating 4 mm internal alignmentmarkers 871 and 825, and then mating 3 mm external alignment markers 872and 824. Once the antenna holder 87 and chassis 82 are completely mated,the needle holder 84 and the triggering arm 83 must be loaded into theantenna holder from the rear end using grade 416 Stainless steel bowelpins attached to the link. After loading the needle, thesensor-transmitter (also known as antenna) assembly 21 must slidethrough the top end of the previous assembly and then into the externalbody 81.

The hypodermic needle 86 is cylindrical needle that is laser fabricatedto create a rectangular slot through the needle and is re-siliconizedusing 3M MED 4159 to ensure the sensor is not tampered during thepush-retract mechanism. The applicator was prototyped by tough 1500photopolymer resin produced by Formlabs. This applicator can bemanufactured at a low cost by high volume manufacturing like injectionmolding process. An example embodiment of plastic for this process willbe polypropylene as it is resistant to steam sterilization and has goodmechanical strength that enables it to withstand some force required forinsertion process.

The user places the armed applicator on the skin (e.g. upper arm) andgently presses it against the skin. This pushes the needle and theneedle assembly towards the skin. Once the assembly reaches to thetrigger height defined by the groves 823 in the triggering guide 82, thetriggering arm is pushed up by the spring to take the system to theunarmed stage, i.e., the needle is fully retracted inside the system asshown in FIG. 46 . The trigger stage is shown in FIG. 45 and the release(unarmed) stage is shown in FIG. 46 . The trigger stage is important asit enables the release stage after which the applicator operation iscompleted. A close-up of the action of the triggering arm to enable thistrigger operation is shown in FIG. 48 .

Once the applicator assembly is released, the transmitter 2 is leftattached to the external surface of the skin with help of adhesives(biocompatible adhesives like 3M 4077) while the sensor 1 is left underthe skin and the connector 3 is connecting these two. All the steps ofsensor insertion using the applicator are shown in FIG. 43 . An exampleof the transmitter above the skin and the sensor under the skin, afterthe applicator use, is shown in FIG. 47 .

At the end of insertion, the needle is retracted back inside theapplicator to prevent injury or misuse.

The applicator requires small amount of force for operation as theneedle used is small and sharp. The force range depends upon the needlesize and shape and the force and speed with which the user applies thedevice on the skin.

The applicator is unique as it inserts a sensor 1 under the skin that isalready electrically connected to its transmitter 2 via connector 3; thetransmitter 2 and the connector 3 being connected with the connector 220(or mated connection of connectors 32 and 34). This is different thanother wired CGMs (e.g., Medtronic Guardian 3, Dexcom G6, Abbott Libre 2)in which the sensor and the transmitter are initially not attached butrather are attached during the sensor insertion or afterwards. In theIMS case, this is enabled by assembling applicator 8 with thesensor-transmitter assembly 21 in advance. The close-up of this sensorassembly with the applicator is shown in FIG. 49 . It shows that theneedle is not a complete cylinder but has a slot or cut out. In FIG. 49, the slot extends along the entire length of the needle. As such, theneedle resembles an open channel, with its internal space (surrounded bya sidewall of the needle) exposed to exterior by the slot. In otherwords, the slot extends through the side wall of the needle. The purposeof the slot is to enable the needle to retract, leaving the sensor underthe skin. This slot basically allows the needle to move around thesensor (up and down) to enable the needle to pierce the skin going downand also enabling it to go around the sensor during retraction withouttouching the sensor such that the sensor remain under the skin while theneedle retracts. The design can use a hypodermic needle with such aslot. The slot can be either cut into a standard hypodermic needle(e.g., using laser cutting) or the needles can be made using the slot(e.g., by using a tube with a slot). Alternatively, the needles can bemade using a sheet metal process e.g., by using lithographic processing(e.g., wet etching) to create sharp tips followed by forming andstamping. It would be appreciated that the slot may extend alongmajority of the length of the needle (not necessarily the entire lengthof the needle).

The applicator has been used for successful insertion of multiplesensors under the skin of multiple animals and human subjects. Anexample of the complete applicator loaded with a sensor-transmitterassembly and a picture of the successful insertion on an arm is shown inFIG. 50A and FIG. 50B.

Sensor Removal

At the end of sensor life, or when desired, the sensor(s) can beextracted by pulling on the external transmitter which is connected tothe sensor using the two-wire flexible connector. The change in sensorreading (e.g., temperature sensor, electrochemical sensors) is used toconfirm safe removal of the device.

Example 1

The CMOS sensor circuits were designed in CAD tools using process designkits (e.g., TSMC 180 nm PDK) and were sent to a CMOS foundry (TSMC) forfabrication. After the fabricated sensors were received, those wereinspected to match the dimensions and similar physical features with thesubmitted design. Afterwards, postprocessing was started to replace thetop metal with more suitable metals. Briefly, AZ5214E resist was spun at4000 rpm, baked at 95 degrees C. for 5 minutes, and exposed using i-Line(e.g., 365 nm UV radiation) exposure in a mask aligner (e.g., MA6) for 5seconds. Next, a post-exposure bake at 120 degrees C. for 5 minutes wasperformed followed by a flood exposure for 3 seconds. Next, the resistwas developed in AZ300 developer. This was followed by sputtering of Ti(e.g., 20 nm) followed by Pt (100 nm). After sputtering, a conformalcoating is achieved. The next step of post-processing was lift-off toremove metal layers from the unwanted regions by soaking the coateddevices in Acetone followed by agitation in an ultrasonic bath. Next,the dies were cut using mechanical saw dicing to cingulate the multiplesensors from one design. Next, several sensors (9 in one example) weredie attached to a flexible PCB substrate using 31CL epoxy, followed bybaking at 100 C for 1 hour to cure the epoxy. Next, the sensor pads werewire-bonded to corresponding pads on the flexible PCB using 1 mil goldwire in a K&S wire bonder. The wire bonds were then encapsulated in aninsulating material (e.g., 31CL) which was then cured at 100 C for 1hour. Afterwards, a connector was soldered to the other end of theflexible PCB to form an interface to the transmitter. Afterwards, theenzyme was immobilized on the sensor in a hydrogel (e.g., a cross-linkedprotein matrix) at a thickness of 3 μm. This was done throughimmobilization of the enzyme GOx (Glucose Oxidase) in a hydrogel createdby Human Serum Albumin (HSA) with glutaraldehyde as the crosslinkingagent. The dispensed solutions were made by mixing GOx and HSA (1200 mg,and 1000 mg respectively) in 15 ml DPBS and a crosslinking agentsolution of 1 w/w glutaraldehyde in DPBS. A 1 microliter solution wasdispensed on the sensor, followed by spinning at 1000 rpm to control thethickness of the hydrogel layer more precisely. Next, a polymer layer(e.g., 2.5% PU in THF with HMDI, Jeffamine, ED-600, DMS-A15 Mn 3000,DEG, Dibutyltin bis(2-ethylhexanoiate) as described in patent 4) on theenzyme layer. It serves to control glucose and oxygen diffusion tooptimize the sensor response. The chemistry stack was allowed to dry ina convention oven at 35° C. for 12 hours. Then the sensors are allowedto stabilize in PBS (Sigma Aldrich, St. Louis MO) for 24 hours andcharacterized (sample testing) for glucose response. Next, the flexiblePCB panel was laser cut to separate individual sensors for next steps.

After cutting, the sensors from the panel were tested in-vitro. First, asample sensor was powered up by connecting it to the transmitter andinitial current was allowed to stabilize. Once a stable baseline wasachieved, the test solution was spiked with small volumes of a stockglucose solution to create an increasing glucose concentration in thesolution. The concentration was measured using a benchtop glucosemonitor from Yellowstone Instruments (YSI 2700) for validation. Threereadings were taken from YSI for statistical validation. As the solutionwas spiked with more and more glucose, the glucose concentration in thetest solution was increased sequentially. The testing results showedthat the sensor current increased when glucose concentration wasincreased in a statistically significant manner. Similar results wereobtained from several sensors to establish the validity of the sensor.

FIG. 51 shows testing results of the sensor assembly testing. Thesolid-state sensors were tested in different concentrations of hydrogenperoxide as it is the most common analyte generated by the oxidase-basedenzymes in the presence of their substrate e.g., Glucose Oxidase in thepresence of glucose, Lactate Oxidase in the presence of Lactate. Itshows that the sensor is quite sensitive to hydrogen peroxide andgenerates a significantly high current (in 10's of nanoamperes)proportional to peroxide concentration. It shows that the 2× gaincircuit design shows 2× the current of the 1× gain design, as desired.It also shows the repeatability of the same sensor tested in thesolution of the same concentration several times in FIG. 51B. FIG. 51Cshows that for the integrated sensor having 3 district sensors on-chip,all 3 are sensitive to peroxide with small variation among them due toprocess variations (e.g., surface cleaning after packaging). Thesevariations can be substantially decreased by increasing process control,by using trimming, and by adding algorithmic features (e.g., calibratingthe individual sensors to different slope factors to minimize thedifferences in current readings). This scheme allows for redundantsensing of and calibration for the same analyte (e.g., glucose), sensingof different analytes (e.g., glucose, BHB, creatinine, urea, etc.) onthe same chip.

FIG. 55 shows the response of a single sensor to increasing glucoseconcentrations after surface coating with glucose oxidase and PUmembrane. The nonlinearity in response can be improved by adding thethickness of the PU membrane. FIG. 55 shows the response of anothersingle sensor to increasing glucose concentrations after surfacefunctionalization with glucose oxidase and PU membrane. Both these are1× gain sensors and show very similar responses to the same glucoseconcentrations. FIG. 55 shows the response of a sensor with 2× gain toglucose. It shows almost 2× the current as compared to the 1× gainsensors, thus proving this design feature. FIG. 55 shows the result oftesting a glucose sensing platform with 3 on-chip glucose sensors. Theplatform was tested in PBS in different glucose concentrations as shownon the graph, and the current for all 3 sensors was read using thetransmitter 2 and the reader 4. It shows that there are variations amongsensors for the same glucose concentration, likely due to the processvariations. These variations can be decreased by improving processcontrol (e.g., surface cleanliness). In any case, the results indicatethat if the median of the 3 readings is used, it minimizes the effect ofsensor-to-sensor variations and provides a more accurate glucose readingas compared to the reading from any single sensor.

FIG. 52A, FIG. 52B, and FIG. 52C show a similar test for a 3-sensorplatform, showing time variations in sensor current as it is subjectedto varying glucose concentrations. It shows the 3 on-chip sensorsrespond in a synchronized manner to glucose, with small variations inactual current due to process variations. The temporary spikes are dueto spiking the test solution with glucose. These don't appear in thebody as there is no spiking of the body fluids with concentratedglucose.

FIG. 53A, FIG. 53B, and FIG. 53C show a similar test for a 3-sensorplatform, showing time variations in sensor current as it is subjectedto varying glucose concentrations. It shows the 3 on-chip sensorsrespond in a synchronized manner to glucose, with small variations inactual current due to process variations. This shows that multiplecopies of the platform perform similarly within the bounds of processvariations.

FIG. 54A and FIG. 54B show the advantage of average and median of the 3on-chip sensors' readings on the same device to decrease the variationbetween 3 on-chip sensors.

Example 2

The sensors here were made using the methods described in example 1. Inaddition, some of the sensors were coated with an additional layer ofPVA using methods described in example 3.

FIG. 56 shows the results of testing different chemistry layers on testsensors. It shows that the sensor coated with the stack of GlucoseOxidase hydrogel, Polyurethane, and Polyvinyl Alcohol (GOx+PU+PVA) hassmaller current than the sensor coated with just GOx+PVA. The GOxcoating was done with spin coating GOx+HSA+Glutarladehyde mixture, thePU coating was done by spin coating PU mixture in THF, and PVA coatingwas done by spin coating PVA mixture in DI water, as detailed in example4. The mixtures were all prepared as described below (example 3). Thelesser sensor current with PU layer as compared to without PU layer isunderstandable as the PU layer regulates the diffusion of glucose andhence decrease the glucose generated current. These results do indicatethe suitability of both chemistry stacks for further testing. Thedevices with GOx+PU+PVA were selected for human testing as betterlinearity can be achieved from this chemistry stack over a largerglucose range as required in later testing in diabetic patients.

Example 3

The solid-state sensors for this example were prepared same as inExample 2. After postprocessing, the devices were coated with GOx in ahydrogel (e.g., a cross-linked protein matrix) at a thickness of ˜3 μm.This was done through immobilization of the enzyme GOx (Glucose Oxidase)in a hydrogel created by Human Serum Albumin (HSA) with glutaraldehydeas the crosslinking agent. The dispensed solutions were made by mixingGOx and HSA (1200 mg, and 1000 mg respectively) in 15 ml DPBS and acrosslinking agent solution of 1% w/w glutaraldehyde in DPBS. A 1microliter solution was dispensed on the sensor, followed by spinning at1000 rpm to control the thickness of the hydrogel layer more precisely.Next, a polymer layer (e.g., 2.5% PU in THF with HMDI, Jeffamine,ED-600, DMS-A15 Mn 3000, DEG, Dibutyltin bis(2-ethylhexanoiate) asdescribed in patent 4) on the enzyme layer. It serves to control glucoseand oxygen diffusion to optimize the sensor response. Finally, a layerof 4% PVA solution was also spun-coated on the PU layer. Namely, PVA gelsolution was formed by adding 0.02% (0.005%-0.1% acceptable)Glutaraldehyde to 4% PVA solution (v/v), vortexing to mix. Surface ofthe PU coated sensors were activated with plasma treatment (80-250 W,5-90 seconds, Oxygen plasma is preferred). 2 ul of gel solution was thendeposited on PU coated sensors and spun at 500 rpm.

The chemistry stack was then baked for 12 hours (2-48 hours acceptable)at 40 degrees Celsius (30-42 C acceptable), in an incubator.

Then the sensors were allowed to stabilize in PBS (Sigma Aldrich, St.Louis MO) for 24 hours and characterized (sample testing) for glucoseresponse. Next, the flexible PCB panel was laser cut to separateindividual sensors for next steps.

After cutting, the sensor was soaked in PBS for 2 hours (overnight ispreferable) before testing. Next, the sensors were tested in differentglucose concentrations and at different temperatures (close to bodytemperature) to determine their suitability for sensing in the body.

The validated sensing devices were then connected to the transmitter andthe assembly was loaded inside the applicator and the entire assemblywas sterilized (e.g., by Synergy Health (San Diego, CA)) usingelectron-beam sterilization using 25 kGray of e-beam irradiation. Thesesensors can be placed in the upper arm using the applicator (aka theinjector). After the sensors are placed in the skin, their data is readvia the IMS transmitter and the IMS reader application on a PC.

FIG. 57A shows the modeling and laboratory results of the IMStemperature sensor testing. FIG. 57B shows the sensor current shoulddecrease with increase in temperature as per the simulation.

FIG. 58A shows the sensor can track environmental temperature changesfor more than 1 week when kept inside a PBS solution without anytemperature control.

FIG. 58B shows the temperature sensor can detect if the sensor is at theroom temperature (before insertion) or at the body temperature (afterinsertion in a person). Hence, it can be used as an additionalvalidation check of successful sensor insertion. This is in addition tothe change in sensor current in air (before insertion) and in the tissue(after insertion). Such multi-modal validation of insertion can be morereliable than a single sensor validation.

Example 4

The solid-state sensors for this example were prepared same as inexample 3. After postprocessing, the devices were coated with GOx in ahydrogel (e.g., a cross-linked protein matrix) at a thickness of −3 μm.This was done through immobilization of the enzyme GOx (Glucose Oxidase)in a hydrogel created by Human Serum Albumin (HSA) with glutaraldehydeas the crosslinking agent. The dispensed solutions were made by mixingGOx and HSA (1200 mg, and 1000 mg respectively) in 15 ml DPBS and acrosslinking agent solution of 1% w/w glutaraldehyde in DPBS. A 1microliter solution was dispensed on the sensor, followed by spinning at1000 rpm to control the thickness of the hydrogel layer more precisely.Next, a polymer layer (e.g., 2.5% PU in THF with HMDI, Jeffamine,ED-600, DMS-A15 Mn 3000, DEG, Dibutyltin bis(2-ethylhexanoiate) asdescribed in patent 4) on the enzyme layer. It serves to control glucoseand oxygen diffusion to optimize the sensor response. Finally, a layerof 4% PVA solution was also spun-coated on the PU layer. Namely, PVA gelsolution was formed by adding 0.02% (0.005%-0.1% acceptable)Glutaraldehyde to 4% PVA solution (v/v), vortexing to mix. Surface ofthe PU coated sensors were activated with plasma treatment (80-250 W,5-90 seconds, Oxygen plasma is preferred). 2 microliter of gel solutionwas then deposited on PU coated sensors and spun at 500 rpm.

The chemistry stack was then baked for 12 hours (2-48 hours acceptable)at 40 degrees Celsius (30-42 C acceptable), in an incubator.

Then the sensors were allowed to stabilize in PBS (Sigma Aldrich, St.Louis MO) for 24 hours and characterized (sample testing) for glucoseresponse. Next, the flexible PCB panel was laser cut to separateindividual sensors for next steps.

After cutting, the sensor was soaked in PBS for 2 hours (overnight ispreferable) before testing. Next, the sensors were tested in differentglucose concentrations and at different temperatures (close to bodytemperature) to determine their suitability for sensing in the body.

The validated sensing devices were then connected to the transmitter andthe assembly was loaded inside the applicator and the entire assemblywas sterilized (e.g., by Synergy Health (San Diego, CA)) usingelectron-beam sterilization using 25 kGray of e-beam irradiation. Thesesensors can be placed in the upper arm using the applicator (aka theinjector). After the sensors are placed in the skin, their data is readvia the IMS transmitter and the IMS reader application on a PC.

FIG. 60A, FIG. 60B, and FIG. 60C show the response of two IMS sensorsprepared using the method above and inserted in a human subject, incomparison with a contour blood glucose meter, a Dexcom G6 comparator(inserted a day before the IMS sensor for FIG. 60A; was not availablefor use for experiment shown in FIG. 60B), and an abbot libre sensor(inserted the same day as the IMS sensor). These results show that theIMS sensor can track glucose changes created by mealtime excursion. Thehuman subject was prediabetic and was told to fast overnight. Thisresulted in their starting glucose level to be close to 100 mg/dl. Thenthey were given a controlled meal which increases their blood glucoselevel (excursion), causing all the sensors to pick-up the increase. Thestudy could only capture 1 hour of this excursion as it was limited induration as per IRB approval and the user had to leave early to attendan event. The user returned after 3 days to get the sensors removed. Noadverse safety event was observed. After the first sensor was removed, asecond study was conducted with the user via a different sensor. Thepurpose this time was to study the effect of low glucose on sensorperformance. The user did start around 100 mg/dl and went to around 82mg/dl during the duration of the study. The IMS sensor (as shown in FIG.60B) could detect the decrease in glucose in a smooth manner whereas theLibre 2 struggled in reading the low glucose value accurately.

FIGS. 61A and 61B show the IMS sensor response before and after thehuman testing. FIG. 61A shows the sensor had sufficient linearity andsensitivity before inserting it in a person. FIG. 61B shows the sensorwas still functioning after it was removed from the human subject. Thisshows that the sensor can function for longer duration (the current testwas subject to IRB approval duration of the study). These resultsindicate the suitability of the IMS CGM platform, including itsconstruction and chemistry, for glucose monitoring in humans.

The study also showed the value of temperature sensor in detectingsensor insertion in the body. FIG. 58B shows the sensor can detect achange from room temperature (around 25° C. in this example) to bodytemperature (around 37.5° C. in this example). This sudden and largetemperature change and stabilization near the body temperature indicatesthe sensor insertion in the body. It also shows the change in tissuetemperature, as detected by the sensor, due to variations in thesurrounding temperature are typically smaller and can also be detectedby the temperature sensor.

The temperature sensor can also detect changes in the peripheral bodytemperature due to an impending hypoglycemia, thus leading to theplatform's ability for a multi-modal (e.g., electrochemical and thermalschemes or modes in this example) hypoglycemia detection which would bebetter than just relying on one measurement of glucose to make thisdecision. FIG. 59A and FIG. 59B show the effect of decreasing glucoseconcentration (going towards hypoglycemia) on the peripheral (e.g.,subcutaneous tissue) temperature. The decrease in temperature wasmeasured by both sensors on the same person (temperature sensor readshigher for lower temperature and reads lower for higher temperature).This is consistent with some studies which find out a change inperipheral temperature in hypoglycemia. Such multi-modal hypoglycemiadetection can provide early warning as well as increase the reliabilityand accuracy of hypoglycemia detection, a key limitation of current CGMplatforms.

An increase in environment temperature can also cause an increase insubcutaneous tissue temperature which can increase the enzymaticactivity which in turn can increase sensor current. This can makedetecting hypoglycemia harder in a hot environment. This can bemitigated by using the IMS temperature sensor which can calibrate thesensor response according to the surrounding temperature and hence canmitigate this risk substantially as compared to other sensors withoutsuch temperature sensor. An opposite effect (considering higher glucoseas hypoglycemia) can occur in colder environments. The integratedtemperature sensor can avoid this problem as well.

Example 5

The IMS sensors for this example were prepared using the same methods asdescribed in Example 4.

This example shows that the IMS sensor can track a reference CGM (DexcomG6 in this case, inserted 1 day before IMS sensor insertion to enable itto get past day 1 issues) over a period of 3 days (limited by testduration, not sensor lifetime). The Day 1, Day 2, and Day 3 graphs ofFIG. 62A, FIGS. 62B, and 63C show that the IMS CGM platform has similartrends as Dexcom G6 (considered to be the most accurate CGM in themarket), whereas the Libre 2 CGM reads lower than the two sensors. TheDay 1 graph of FIG. 62A shows that the IMS sensor (Day 1 of IM sensor)is well matched to the Dexcom G6 data (Day 2 for Dexcom G5 since it wasinserted a day before the IMS sensor to avoid day 1 noise issues presentin current CGMs). Libre 2 has similar trends but shows lower glucoseconcentrations for some reason. Also, there are some locations (e.g.,18:43) where the Dexcom and the IMS sensor show different trend than theLibre 2. Similar trend can be seen on Day 2 in FIG. 62B. There aresometimes (e.g., 6:36, around noon) where the IMS sensor matches betterwith the Libre 2 trend than the Dexcom G6 trend. On Day 3 in FIG. 62C,the test subject's glucose goes through a nice and clear excursion cyclewhich is traced by the IMS sensor very well except towards the end whenthe sensor support (skin tape) starts peeling off due to the end of thewear period for the tape. Overall, IMS values tracked the Dexcom G6 dataquite well while Libre 2 data was significantly off from the Dexcom G6reading. There was no contour reference data available beyond first fewhours on day 1 as the test subject didn't want to prick routinelyoutside the clinic. This IMS sensor data only uses signal averaging tomatch the IMS data points with the reference method and doesn't rely onany detailed signal processing methods described earlier. This showsthat the IMS sensor is fundamentally accurate by design and doesn'trequire complex signal processing to improve its output. Signalprocessing can be used though to improve certain features likeprediction of future values and detection of hypoglycemia before itcauses health issues or fainting.

The three-day wear period was limited by the transmitter tape peel off.The sensor was still operational after the test (as shown in FIG. 61B)which shows the device can perform for longer duration. We have sincethen found a better tape that causes less skin irritation and worksbetter with the skin for a longer duration.

Example 6

The IMS sensors for this example were prepares using the same methods asdescribed in Example 4. This example shows the value of having multipleon-chip sensors for glucose sensing. FIG. 63 shows that in case somesensors (e.g., sensor 3) run into a noise issue momentarily, othersensors can be used to detect this issue (e.g., by detecting theerroneous sensor's reading to be >10% different than other sensors).After this, this erroneous sensor's weight can be changed (e.g., 0 todiscard its reading fully for some time) to increase the accuracy of theoverall reading (e.g., weighted average of the 3 sensors). The algorithmto perform these calculations is provided in the “multi-sensor signalprocessing schemes” section. FIG. 63 is plotted based upon Table 1. Theresults show that the error in the weighted average (after the ‘outlier’sensor readings are discarded, as shown by black dotted line in FIG. 63) is <2% whereas the error in the direct (simple) arithmetic average(assuming equal weights for all sensors at all times; as shown by thegreen line in FIG. 63 ) is >15%.

This example shows the value of multi-sensor scheme. In case of a singlesensor system, a sensor error is difficult to be detected and even if itcan be detected (e.g., based upon sudden changes being larger thanphysiologically possible), it is difficult to detect the actual value.In case of the IMS design having multiple sensors, this is not a problemas a single sensor reading erroneous readings can be detected bycomparing its reading with others and hence its weight from the weightedaverage can be reduced or its reading can be eliminated from thecalculations to minimize its effect. This makes the IMS platform moreaccurate and reliable as compared to others. Also, once several readingsfrom a sensor have been obtained, the next reading can be predictedbased upon rate of change and physiological models. Hence, the sensorreadings can be compared against this expected reading to determineerroneous readings and those can be eliminated from the calculations(this is not shown in this example for simplicity reasons but can bedone).

Example 7

The IMS sensor for this example was prepares using the same methods asdescribed in Example 4.

The example shows the rate of change of the Libre 2, IMS, and Dexcom G5sensors in response to a step change in glucose concentration as theresult of a spike (addition of high-concentration stock solution to thetest buffer under stirring). The reference concentration was measuredusing a YSI 2700 at regular intervals. Please see FIG. 64 . The exampleshows that Libre 2 couldn't detect the lower concentration (90 mg/dl)and kept reading ‘glucose concentration too low’ till the higherconcentration (180 mg/dl) was achieved in the test solution. It stilltook around 20 minutes after the spike to register the change and wasoff from the YSI reference reading by a large margin (>120 mg/dl) forthe rest of the experiment. Dexcom G5 performed better and was readingreasonably accurate before the spike (around 100 mg/dl). After thespike, it did detect the change initially within 5 minutes but took >20minutes in total to get close to the actual value as measured by theYSI. The IMS sensor was reading accurately before the solution wasspiked. After the spike, it took <5 minutes (approximately 4 minutes) toreach to the steady-state value close to the actual value. Therefore,according to FIG. 64 , the IMS sensor is capable of detecting a rate ofchange of glucose concentration at approximately 22.5 mg/dl/min. This isdue to the thin diffusion barriers (e.g., thin enzyme layer, thinpolymer membrane, thin biocompatibility layer as described above) ascompared to the commercially available devices (which use thickerdiffusion barriers to minimize the effect of larger manufacturingvariations), enabled by the planar nature of the IMS device (e.g.,enable spin coating).

To allow the IMS sensor to detect a rate of change of >10 mg/dl/minute,the thickness of the enzyme layer is preferably less than 3500 nm, andmore preferably, less than 1000 nm. In an example, the enzyme layer maybe between 200 nm and 800 nm in thickness, and more specifically,between 600 nm and 800 nm in thickness.

To allow the IMS sensor to detect a rate of change of >10 mg/dl/minute,the polymer membrane preferably has a thickness of between 200 nm and10500 nm, and more preferably, has a thickness of between 200 nm and1500 nm. The thickness of the biocompatibility layer may be between 1000nm and 20000 nm. A biocompatibility layer with a thickness of less than1000 nm may be used alternatively.

The faster response is also enabled by the fact that the IMS sensor isvery accurate and hence doesn't need to perform heavy signal processing(e.g., long moving averages) to reduce the effect of sensor noise.Hence, the IMS sensor has the least lag among the 3 CGMs tested in thisexperiment. This effect is expected to translate to in-vivo sensing aswell since the transport mechanisms through the surface chemistry layersare same in both cases. This also shows that the IMS sensor can achievehighest rates of change (e.g., >20 mg/dl/minute) as compared to DexcomG5 (<4 mg/dl/minute) and Abbott Libre 2 (<2 mg/dl/minute).

While specific embodiments of the invention have been described above,it will be appreciated that the invention may be practiced otherwisethan as described and that that the described embodiments are for allpurposes exemplary, not limiting. Various modifications can be made tothe described embodiments without departing from the scope of thepresent invention which is defined by the appended claims.

Further implementations are summarised in the following examples:

-   -   Example 1: A transdermal analyte concentration measurement        system comprising:    -   1) an implantable monolithic integrated circuit containing:        -   an integrated sensing element which senses one or more            analytes and generates a signal representative of an analyte            concentration;        -   an integrated sensor signal acquisition unit which receives            and processes the signal from the integrated sensing            element;        -   an integrated communication unit connected to the integrated            sensor signal acquisition unit which transmits data            representative of said analyte concentration; and        -   an integrated power management unit connected to and            providing power to the sensing element, the sensor signal            acquisition unit, and the communications unit; and    -   2) a transmitter, configured for operation outside of the body        of a user, connected by a wire to the integrated communication        unit and integrated power management unit of the implantable        monolithic integrated circuit, said transmitter providing power        to the integrated power management unit of the implantable        monolithic integrated circuit via the wire and receiving data        from the integrated communication unit of the implantable        monolithic integrated circuit via the wire, and said wire        configured to cross from the skin surface into the user to span        the dermis of a user.    -   Example 2: The transdermal analyte concentration measurement        system of example 1, wherein the integrated sensing element of        the implantable monolithic integrated circuit includes one or        more integrated electrodes.    -   Example 3: The transdermal analyte concentration measurement        system of example 2, wherein at least one of the one or more        integrated electrode comprises a conductive surface of one or        more conductive materials.    -   Example 4: The transdermal analyte concentration measurement        system of example 2 or 3, wherein the integrated electrodes are        made using lithography.    -   Example 5: The transdermal analyte concentration measurement        system of any one of examples 2 to 4, wherein the integrated        electrodes are coated with a hydrogel including a cross linking        agent, an enzyme, and a proteinaceous material.    -   Example 6: The transdermal analyte concentration measurement        system of example 5, wherein the hydrogel further comprises a        co-protein.    -   Example 7: The transdermal analyte concentration measurement        system of example 5 or 6, wherein: the cross-linking agent        comprises glutaraldehyde; the enzyme comprises glucose oxidase;        and the proteinaceous material comprises human serum albumin.    -   Example 8: The transdermal analyte concentration measurement        system of any one of example 5 to 7, wherein the hydrogel        further comprises a co-protein of catalase or horseradish        peroxidase.    -   Example 9: The transdermal analyte concentration measurement        system of any one of examples 1 to 8, wherein the wire is        comprised within a flexible printed circuit board.    -   Example 10: A glucose sensor system comprising:    -   a transmitter containing a battery for placement on top of        patient skin;    -   a transcutaneous connector comprising at least one conductive        path; and    -   a potentiostat and an electrochemical sensing element for        placement beneath the patient skin;    -   wherein the potentiostat is electrically coupled to the        transmitter via the transcutaneous connector, and the        electrochemical sensing element is configured to sense glucose        concentration and generate an electrical signal representative        of the glucose concentration, and wherein the potentiostat is        electrically connected to the electrochemical sensing element.    -   Example 11: The glucose sensor system of example 10, wherein the        potentiostat is continuously powered by the battery.    -   Example 12: The glucose sensor system of example 10 or 11,        wherein the glucose sensor system is a continuous glucose sensor        system.    -   Example 13: The glucose sensor system of any one of examples 10        to 12, wherein the transcutaneous connector is a flexible        connector.    -   Example 14: The glucose sensor system of any one of examples 10        to 13, wherein the potentiostat is placed at a depth of 1 to 10        mm beneath the patient skin.    -   Example 15: The glucose sensor system of any one of examples 10        to 14, wherein the electrochemical sensing element comprises at        least one working electrode coated with a chemistry which        converts glucose concentration into current.    -   Example 16: The glucose sensor system of example 15, wherein the        potentiostat is connected within half a millimeter to the        entirety of at least one working electrode.    -   Example 17: The glucose sensor system of example 15 or 16,        wherein the electrodes are patterned to increase surface area.    -   Example 18: The glucose sensor system of any one of examples 15        to 17, wherein the electrodes are patterned by forming pillars.    -   Example 19: The glucose sensor system of example 18, wherein the        pillar spacing is 0.25 μm-25 μm and height 0.1 μm-10 μm.    -   Example 20: The glucose sensor system of any one of examples 15        to 19, wherein the electrodes are integrated with the        potentiostat into the same CMOS die.    -   Example 21: The glucose sensor system of example 20, wherein the        CMOS die is from 30 microns to 600 microns in thickness, 500        microns to 10,000 microns in length and in a range from 100        microns to 4,000 microns in width.    -   Example 22: The glucose sensor system of example 20, wherein the        CMOS dies is from 50 microns to 150 microns in thickness, 1,500        microns to 3,000 microns in length and in a range from 100        microns to 4,000 microns in width.    -   Example 23: The glucose sensor system of any one of examples 10        to 22, wherein the potentiostat is at a depth of 1 to 5 mm        beneath the patient skin.    -   Example 24: The glucose sensor system of any one of examples 10        to 23, wherein the electrochemical sensing element comprises at        least three working electrodes with each connected to a        respective potentiostat.    -   Example 25: The glucose sensor system of any one of examples        10-24, further comprising an analog to digital converter        connected to the potentiostat.    -   Example 26: The glucose sensor system of example 25 as dependent        from example 24, further comprising a control logic programmed        to process digital information from the analog to digital        converter and consider the information from one or more of the        at least three working electrodes.    -   Example 27: The glucose sensor system of example 26, wherein the        control logic is configured to denoise the information from one        or more of the at least three working electrodes.    -   Example 28: The glucose sensor system of any one of examples 10        to 27, wherein the transcutaneous connector comprises a printed        circuit board.    -   Example 29: The glucose sensor system of examples 10 to 28,        further comprising a temperature sensor monolithically        integrated with the potentiostat.    -   Example 30: The glucose sensor system of example 29, wherein the        temperature sensor comprising a bandgap circuitry to generate        current proportional to absolute temperature (PTAT) and current        complementary to absolute temperature (CTAT).    -   Example 31: The glucose sensor system of example 29 or 30 as        dependent from example 15, further comprising a control logic        programmed to process information from the at least one working        electrode by taking into account information from the        temperature sensor.    -   Example 32: The glucose sensor system of any one of examples 10        to 31, wherein the transcutaneous connector comprises at least        two conductive paths.    -   Example 33: The glucose sensor system of example 32, wherein one        of the conductive paths is a ground connected to the        potentiostat and another one of the conductive paths functions        as data over power.    -   Example 34: The glucose sensor system of examples 10 to 33,        wherein the potentiostat is formed in a CMOS die.    -   Example 35: The glucose sensor system of example 34, wherein the        CMOS die is bonded to the flexible connector and the at least        one conductive path by wire bonding.    -   Example 36: The glucose sensor system of example 35, wherein the        CMOS die is bonded to the flexible connector and the at least        one conductive path by flip chip packaging.    -   Example 37: The glucose sensor system of examples 16-36 as        dependent from example 15, wherein the at least one working        electrode is coated with a glucose oxidase hydrogel.

The invention claimed is:
 1. A glucose sensor system comprising: atransmitter for containing a battery, the transmitter being forplacement on top of patient skin; a transcutaneous connector comprisingat least one conductive path; and an implantable monolithic integratedcircuit for placement beneath the patient skin, wherein the implantablemonolithic integrated circuit comprises a potentiostat and anelectrochemical element, said electrochemical sensing element comprisingat least one working electrode patterned to increase surface area andcoated with a chemistry which converts glucose concentration intocurrent; wherein the potentiostat is electrically coupled to thetransmitter via the transcutaneous connector, and the electrochemicalsensing element is configured to sense glucose concentration andgenerate an electric signal representative of the glucose concentration,and wherein the potentiostat is electrically connected to theelectrochemical sensing element.
 2. The continuous analyte monitoringsystem of claim 1, wherein the potentiostat is connected within half amillimeter to the entirety of at least one working electrode.
 3. Thecontinuous analyte monitoring system of claim 1, wherein the at leastone working electrode is patterned by forming holes.